Biodegradable, Thermally Responsive Injectable Hydrogel for Treatment of Ischemic Cardiomyopathy

ABSTRACT

Provided are novel biocompatible copolymers, compositions comprising the copolymers, and methods of using the copolymers. The copolymers are non-toxic and typically have an LCST below 37° C. Compositions comprising the copolymers can be used for wound treatment, as a cellular growth matrix or niche and for injection into cardiac tissue to repair and mechanically support damaged tissue. The copolymers comprise numerous ester linkages so that the copolymers are erodeable in situ. Degradation products of the copolymers are soluble and non-toxic. The copolymers can be amine-reactive so that they can conjugate with proteins, such as collagen. Active ingredients, such as drugs, can be incorporated into compositions comprising the copolymers.

This application claims priority to U.S. Provisional Patent ApplicationNo. 61/979,244, filed Apr. 14, 2014, the contents of which areincorporated herein by reference in their entirety.

NOTICE OF GOVERNMENT SUPPORT

This invention was made with government support under Grant No. HL105911awarded by the National Institutes of Health. The government has certainrights in the invention.

BACKGROUND

A thermoresponsive, biodegradable elastomeric material is describedherein, along with methods of making the material and uses for thematerial, particularly uses of the material in repairing defects inheart muscle.

Injectable thermally responsive hydrogels with a lower critical solutiontemperature (LCST) below body temperature represent promisingbiomaterials for a variety of biomedical applications, includingregional tissue mechanical support as well as drug and cell deliveryapplications. Generally, the LCST-based phase transition occurs uponwarming in situ as a result of entropically-driven dehydration ofpolymer components, leading to polymer collapse. Various naturallyderived and synthetic polymers exhibiting this behavior have beenutilized. Natural polymers include elastin-like peptides andpolysaccharides derivatives, while notable synthetic polymers includethose based on poly(n-isopropyl acrylamide) (pNIPAAm), and amphiphilicblock copolymers, often containing poly(ethylene glycol). The structureof pNIPAAm, containing both hydrophilic amide bonds and hydrophobicisopropyl groups, leads to a sharp phase transition at the LCST. Studiessuggest that the average number of hydrating water molecules per NIPAAmgroup falls from 11 to ˜2 upon the hydrophobic collapse above the LCST(32° C.)

pNIPAAm based polymers have been extensively studied as injectablebiomaterials for tissue regeneration and drug delivery, yet pNIPAAmitself is a non-biodegradable polymer with a constant LCST ofapproximately 32° C., which prevents ready clearance from the body atphysiologic temperature. This limitation of pNIPAAm has provided themotivation for developing biodegradable NIPAAm-based polymers byconjugating the pNIPAAm with natural biodegradable segments such asMMP-susceptible peptide, gelatin, collagen, hyaluronic acid and dextran.However, these may be only partially bioabsorbable since sufficientlylong pNIPAAm segments would remain non-soluble following removal of thenatural segments.

Copolymers formed from NIPAAm and monomers with degradable side chainscomprise another category of NIPAAm-based bioabsorbable, thermallyresponsive hydrogels. Hydrolytic removal of hydrophobic side chainsincreases the hydrophilicity of the copolymer, raising the LCST abovebody temperature and making the polymer backbone soluble. Due to therelative simplicity of the synthetic process, the most investigatedbiodegradable monomers have been HEMA-based monomers, such as2-hydroxyethyl methacrylate-polylactide (HEMA-PLA), 2-hydroxyethylmethacrylate-polycaprolactone (HEMA-PCL) and 2-hydroxyethylmethacrylate-polytrimethylene carbonate (HEMA-PTMC). However, thebackbone remnant following hydrolysis, HEMA, presents hydroxyethyl sidegroups (—CH₂CH₂—OH), which have a relatively limited effect on remnantpolymer hydrophilicity. In previous studies, such hydrogels have beenfound to be either partially bioabsorbable or completely bioabsorbable,but have required the inclusion of considerably hydrophilic monomerssuch as acrylic acid (AAc) in the hydrogel synthesis.

Progressive remodeling of the left ventricular (LV) architecture occursafter myocardial infarction (MI). While initially required formaintenance of cardiac output, this response ultimately leads to LVdysfunction and heart failure in the absence of a recurrent ischemicevent. Even with current optimal therapy, mortality inend-stage-heart-failure amounts to 20-50% per year. Hearttransplantation is applied as the last therapeutic option for patientswith terminal heart-failure, but requests for organ transplantation faroutstrip the number of donor organs. Therefore, new therapeuticstrategies are urgently needed in order to ameliorate both patientprognosis and quality of life.

Following MI, dilatation of the LV cavity has the effect of increasingLV wall tension, which triggers further dilatation of the LV cavity, andprogression down a spiral of adverse cardiac remodeling towards theadvanced stages of cardiac failure. To restore wall tension, theendoventricular circular patch plasty technique (the Dor procedure) andpartial left ventriculectomy (the Batista procedure) have beenclinically implemented for severe cardiac dilation and dysfunction manyyears after an infarction. Employing a similar strategy to limit theremodeling pathway at an earlier stage, epicardial restraint therapies,such as the Acorn Cardiac Support Device, and the Paracor device havebeen investigated. However, these both apply materials that arenon-biodegradable and result in a permanent foreign body encapsulatingthe epicardium.

Using biodegradable and elastic polyester urethane urea, we recentlyreported that cardiac patch implantation onto a chronic myocardialinfarct prevented further cardiac dilatation and improved contraction,while altering LV wall thickness and compliance. Supported by a finiteelement model simulation, another concept in locally treating thefailing cardiac wall was proposed where a bulking material is injectedinto the infarcted left ventricular wall to positively alter cardiacmechanics and result in a potentially beneficial reduction of elevatedstresses in the infarcted wall. In this numerical model the localsystolic fiber stress distribution was determined in an infarcted LVwall injected with a mechanically passive material. The simulationshowed that injection of a volume 4.5% that of the total LV wall volumeand with a stiffness (elastic modulus) 20% of the natural LV tissue intothe infarct border zone could decrease the fiber stress in the borderzone of the infarct by 20% compared to a control simulation in whichthere was no injection. The mechanical simulation also showed that thisattenuation effect on LV wall stress increased with the injection volumeand the modulus of the injected material.

Thermally responsive hydrogels are particularly attractive materials forinjection therapy following MI since it is possible to inject thenecessary fluid volumes from a syringe maintained below bodytemperature. Upon injection and warming hydrogel mechanical propertiesare increased, the “holding” of the material at the injection site isfacilitated and the mechanical benefit of the injected volume on thecardiac wall is increased.

However, despite the advantages of thermally responsive hydrogels knownto date, many such gels are too robust, specifically in terms of aprolonged time period of hydrolysis and absorption by native tissue. Ifhydrolysis does not occur sufficiently rapidly, infiltration by nativecells can be slowed, which in turn slows the rate of formation of newtissue at the site of injury.

Accordingly, a need exists, and a substantial challenge remains in theart for versatile biocompatible polymer compounds that can serve as cellgrowth substrates, for drug delivery purposes and generally for use inpatients, for example for cardiac remodeling, where thedegradation/hydrolysis time of such compounds can be preciselycontrolled, and where the polymer compounds and their degradationproducts exhibit desirable physical properties.

SUMMARY

Provided herein are compositions comprising thermoresponsive andbiodegradable elastomeric materials; namely copolymers and compositionsand structures, such as hydrogels, comprising the copolymers and methodsof use of those compositions, including a method of treating amyocardial defect, such as an infarct. The copolymers remain fluid belowphysiological temperature (e.g., 37° C. for humans) or at or below roomtemperature (e.g., 25° C.), solidify (into a hydrogel) at physiologicaltemperature, and degrade and dissolve at physiological conditions in atime-dependent manner, which is important for removal of the hydrogelafter an applied surgical or medical procedure.

According to one embodiment, the copolymer comprisesN-isopropylacrylamide (NIPAAm) residues (a residue is the remainder of amonomer incorporated into a polymer), hydroxyethyl methacrylate (HEMA)residues, the polyester macromer methacrylate-polylactide (MAPLA)macromer residues, and methacrylic acid (MAA) monomers. Alternately, thecopolymer comprises NIPAAm residues, acrylic acid (AAc) residues, thepolyester macromer hydroxyethyl methacrylate-poly(trimethylenecarbonate) (HEMA-PTMC) macromer residues, and MAA monomers. Alternativesfor NIPAAm residues include N-alkyl acrylamide residues in which thealkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl.Alternatives for HEMA include (hydroxyl (C₁-C₃)alkyl)methacrylate andother methacrylate substituents that can modulate the LCST of thepolymer. Although the size of the copolymers can vary, in one example,the copolymer has an M_(n) of between 20 kD and 35 kD. In anotherexample, the copolymer has a polydispersity index (PDI, M_(w)/M_(n)) ofbetween 1 and 2.

According to another embodiment, the copolymer comprises NIPAAmresidues, N-vinylpyrrolidone monomers (VP), and MAPLA macromer residues,or NIPAAm residues, N-vinylpyrrolidone monomers (VP), and hydroxyethylmethacrylate-poly(trimethylene carbonate) (HEMA-PTMC) macromer residues.Alternately, the copolymer comprises NIPAAm residues, VP monomers, andthe polyester macromer hydroxyethyl methacrylate-poly(trimethylenecarbonate) (HEMA-PTMC) macromer residues. Alternatives for NIPAAmresidues in this embodiment include N-alkyl acrylamide residues in whichthe alkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl.Alternatives for HEMA include (hydroxyl (C₁-C₃)alkyl)methacrylate andother methacrylate substituents that can modulate the LCST of thepolymer. Although the size of the copolymers can vary, in one example,the copolymer has a M_(n) of between 20 kD and 35 kD. In anotherexample, the copolymer has a polydispersity index (PDI, M_(w)/M_(n)) ofbetween 1 and 2.

In each copolymer, the ratio of the constituents of the polyestermacromer may be varied. For example and without limitation, where thepolyester macromer is a poly(trimethylene carbonate (TMC)-containingmacromer), comprising hydroxyethyl methacrylate residues and varyingnumbers of trimethylene carbonate units/residues. In another embodiment,the polyester macromer is a methacrylate-polylactide macromer comprisingmethacrylate residues and varying numbers of lactide residues. Eachcomponent contributes to the desired physical properties of the hydrogelto enable an injectable material for delivering drugs or chemicals,encapsulating and transplanting cells, and injecting into empty cavitiesfor wounds or tissue repair.

An optional amine-reactive component may be included in the copolymer asdescribed above. The amine-reactive group can be a succinimide group, anoxysuccinimide group or an isocyanate group, such as is produced byincorporation of N-hydroxysuccinimide methacrylate (MANHS) or N-acryloxysuccinimide (NAS) monomers into the copolymer. The amine-reactive groupsbind to amine-containing compounds including biomolecules such ascollagen and/or other bioactive or biocompatible materials or factors.Varying amounts of the amine-reactive component may be used, dependingon the desired density of amine-reactive groups, while maintainingdesirable physical and degradation properties of the resultantcopolymer.

The composition of each component in the hydrogel determines the lowercritical solution temperature (LCST) of the hydrogel. At a temperatureless than the LCST, the hydrogel flows easily and can be injected intothe desired shape. When the temperature is increased above the LCST, thehydrogel solidifies and retains its shape. Once solidified, the hydrogelis highly flexible and relatively strong at physiological temperature.For complete removal of the copolymer, the copolymer includeshydrolytically-cleavable bonds that results in soluble, non-toxicby-products, which results in dissolution of the degraded hydrogel andclearance of the degraded components.

In one embodiment, the copolymer has a lower critical solutiontemperature below 37° C., for example 36° C. or lower, 35° C. or lower,34° C. or lower, or, in another embodiment between 10° C. and 34° C.,including increments and sub-ranges therebetween, and in anotherembodiment, less than 20° C. According to one embodiment, the copolymerhas a lower critical solution temperature above 37° C. after its esterbonds are hydrolyzed.

The polymer comprises a polyester macromer, for example and withoutlimitation, a polyester macromer comprising methacrylate-polylactideresidues. In one embodiment, the ratio of methacrylate and lactideresidues in the polyester macromer is from 1:2 (methacrylate:lactide) to1:8, in another, from 1:1 to 1:10, such as 1:1, 1:2, 1:3, 1:4, 1:5, 1:6,1:7, 1:8, 1:9, and 1:10. In another non-limiting example, the polyestermacromer comprises hydroxyethyl methacrylate and trimethylene carbonateresidues. In one embodiment, the ratio of hydroxyethyl methacrylate andtrimethylene carbonate residues in the polyester macromer ranges from1:1 to 1:10, 1:2 to 1:5 or any increment within those ranges, including1:1, 1:2, 1:3, 1:4, 1:4.2, 1:5, 1:6, 1:7, 1:8, 1:9, and 1:10.Amine-containing biomolecules or other compounds, such as proteins,carbohydrates, glycoproteins, etc. can be conjugated to the copolymerthrough an amine-reactive group, when incorporated into the copolymer.In certain embodiments, collagen, gelatin are suitable compounds, forinstance and without limitation, between 1% wt and 10% wt collagen.

A composition comprising the copolymer described herein also maycomprise an aqueous solvent, for example and without limitation, water,saline and phosphate-buffered saline. The composition also can includean active agent, such as, without limitation, one or more of anantiseptic, an antibiotic, an analgesic, an anesthetic, achemotherapeutic agent, a clotting agent, an anti-inflammatory agent, ametabolite, a cytokine, a chemoattractant, a hormone, a steroid, aprotein and a nucleic acid. In one embodiment, where the compositioncomprises a clotting agent, one example of a clotting agent isdesmopressin. In another embodiment, for use (e.g.) in repair of cardiactissue, the active agents are one or both of bFGF and IGF-1. Abiological material, such as a cell or a virus particle may also beincorporated into the composition.

A method is provided of making a thermosensitive copolymer, for exampleand without limitation, a co-polymer described herein, the methodcomprising co-polymerizing NIPAAm, HEMA, MAPLA, and MAA monomers to makea copolymer. In another embodiment, the method comprises co-polymerizingNIPAAm, VP, and MAPLA to make a copolymer. In another embodiment, themethod comprises co-polymerizing NIPAAm, AAc, HEMAPTMC, and MAAmonomers. The monomers can be co-polymerized by any usefulpolymerization method, for example and without limitation by radicalpolymerization methods, such as free-radical polymerization or livingpolymerization methods, such as atom transfer radical polymerization(ATRP).

A method of treating a muscle defect, such as a myocardial defect, suchas an infarct or traumatic injury, is provided. The method comprisesinjecting a composition comprising a copolymer having an LCST of lessthan 37° C., comprising N-alkyl acrylamide residues in which the alkylis one of methyl, ethyl, propyl, isopropyl and cyclopropyl;hydroxyethylmethacrylate (HEMA) residues; methacrylate-polyactide(MAPLA) macromer residues; and methacrylic acid (MAA) residues into thearea of (within, in contact with or in tissue surrounding) a myocardialdefect, or a defect of any muscle tissue. In another embodiment, themethod comprises injecting a composition comprising a copolymer havingan LCST of less than 37° C., comprising N-alkyl acrylamide residues inwhich the alkyl is one of methyl, ethyl, propyl, isopropyl andcyclopropyl; N-vinylpyrrolidone monomers (VP); andmethacrylate-polyactide (MAPLA) macromer residues into the area of(within, in contact with or in tissue surrounding) a myocardial defect,or a defect of any muscle tissue. The injection of either of the abovegeneral class of copolymers may take place from 1 day to 21 days afterinfarction (inclusive of values between those provided here). In anembodiment, the injection of the composition occurs between 7 and 14days (inclusive of values between those provided here) followinginfarction, so as to maximize functional outcome through tissueremodeling while allowing time for the patient to recover from theinsult.

According to another embodiment a method of growing cells is provided,comprising introducing cells into any copolymer composition describedherein to produce a cell construct and incubating the cell constructunder conditions suitable for growth of the cells. The composition cancomprise cell growth media to facilitate cell growth within thecomposition. The cell construct can be administered to a patient (placedin a patient's body at a desired location), such as a human patient. Inanother embodiment, the composition is administered to a patient withoutcells, but so that the patient's cells migrate into the composition. Thecomposition can be administered by an injection into the desired site,such as cardiac tissue within the patient.

For example, the composition may be injected in or around necrotictissue in the heart. In one embodiment, the composition is injectedapproximately 2 weeks after the patient has a myocardial infarction.This injection may take place from 1 day to 21 days after infarction. Inan embodiment, the injection of the composition occurs between 7 and 14days following infarction, so as to maximize functional outcome throughtissue remodeling while allowing time for the patient to recover fromthe insult. The composition also may include one or more active agents,such as, without limitation, an antiseptic, an analgesic, an anestheticand an antibiotic. To facilitate heart repair, or repair of any tissue,or cell growth in general, the composition may comprise, with or withoutother active agents, one or more of a cytokine, a cell growth ordifferentiation agent and a metabolite, such as one or both of bFGF andIGF-1.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1. Synthetic scheme for HEMAPTMC and the copolymerpoly(NIPAAm-co-AAc-co-HEMAPTMC)

FIG. 2. Synthetic scheme for MAPLA and poly(NIPAAm-co-HEMA-co-MAPLA).

FIG. 3. Synthesis of poly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA) andsubsequent reaction of the MANHS mer with growth factor.

FIG. 4. Synthesis of poly(NIPAAm-co-HEMA-co-MAPLA-co-MAA) (pNHMMj)

FIG. 5. Chemical structure of poly(NIPAAm-co-HEMA-co-MAPLA-co-MAA)(pNHMMj). The arrow shows the position of a hydrophobic to hydrophilicalteration upon side-chain hydrolysis.

FIG. 6A-6B. NMR spectra of poly(NIPAAm-co-HEMA-co-MAPLA-co-MAA) (pNHMMj)showing the MAA peak (12 ppm) present in the composition according tothe present invention (FIG. 6A); MAA weight percentages in copolymers asdetermined by NMR and acid titration at feed ratios of 0.5%, 1%, 2%, 5%,and 10% (FIG. 6B).

FIG. 7. pH of supernatants of pNHMMj hydrogels after gelation (a);Fluorescent emission intensity ratio between 540 nm and 440 nm of pNHMMjhydrogels mixed with LysoSensor pH-sensitive dye, excited at 360 nm (b).A higher ratio reflects a lower pH. Data of pNHMM5 and pNHMM10 were notavailable due to fast degradation. * and # indicate significantdifferences between and within groups, respectively.

FIG. 8. Degradation curves of poly(NIPAAm-co-HEMA-co-MAPLA-co-MAA)(pNHMMj) hydrogels with graded MAA content (a); Time for 50% weight lossderived from data in the previous panel (b).

FIG. 9. Transition temperature of pNHMMj hydrogels. Temperature sweep ofshear modulus (G′) of pNHMMj hydrogels (a); Transition temperaturedependence on pH (b).

FIG. 10. Mechanical properties (Young's modulus) for hydrogels formedfrom copolymer compositions according to embodiments of the presentinvention.

FIG. 11. Viability of rSMCs encapsulated in pNHMMj hydrogels. (a-h) LiverSMCs (red) stained with CellTracker in pNHMMj hydrogels 1 d and 7 dafter encapsulation. Green: Fluorescein labeled hydrogels.

FIG. 12. Percentage of live rSMCs after 1 d and 4 d encapsulation inpNHMMj hydrogels, determined by trypan blue staining.

FIG. 13A-13B. Cytotoxicity of degradation products of pNHMMj hydrogels.rSMCs proliferation 1 d (a-e), 3 d (f-j), 7 d (k-o) after seedingdetermined by live/dead staining; (p) MTS assay of the rSMCs.

FIG. 14. Rat heart injected with MAPLA gel labeled with a small amountof fluorescent monomer (A and B).

FIG. 15. Bright field and fluorescent images of excised rat leg musclesinjected with pNHMMj hydrogels. Left column: excised on the same day ofinjection. Right column: excised 21 d after injection. The white massand green fluorescence indicate the hydrogel.

FIG. 16 Immunohistochemical (row (a,b)) and H&E (row (c)) staining ofrat hindlimb muscle injected with pNHMMj hydrogels 21 d after injection.In row (a) and row (b): Blue: DAPI for nucleus, Green: hydrogels, Red:CD68 for macrophages. Images in row (b) are enlarged from correspondingareas in row (a), as indicated by yellow rectangles (may be rotated). Mindicates muscle and G indicates hydrogel.

FIG. 17. Poly(NIPAAm-co-VP-co-MAPLA) copolymers synthesized from NIPAAm,VP, and MAPLA by free radical polymerization.

FIG. 18. Transition time of poly(NIPAAm-co-VP-co-MAPLA) copolymerhydrogels in 37° C. air.

FIG. 19. Poly(NIPAAm-co-VP-co-MAPLA) copolymer hydrogel degradation.

FIG. 20. Mechanical properties of poly(NIPAAm-co-VP-co-MAPLA) copolymerhydrogels.

FIG. 21. Cytotoxicity of degradation products ofpoly(NIPAAm-co-VP-co-MAPLA) copolymer hydrogels evaluated by live/deadstaining.

FIG. 22. Cytotoxicity of degradation products ofpoly(NIPAAm-co-VP-co-MAPLA) copolymer hydrogels evaluated by MTS assay.

FIG. 23. H&E staining of tissue following injection ofpoly(NIPAAm-co-VP-co-MAPLA) copolymer hydrogels.

FIG. 24. Adeno-associated virus release from poly(NIPAAm-co-VP-co-MAPLA)copolymer hydrogels.

FIG. 25. Change of supernatant pH during degradation of pNHMMj hydrogelsin PBS.

FIG. 26. Weight loss of pNHMM10 incubated in regular PBS (pH 7.4) andbasic PBS (pH 9.5, mediated by NaOH).

FIG. 27. SMCs cultured for another 7 d after retrieving from insidepNHMMj hydrogels and seeding on TCPS.

FIG. 28. Surface and cross section fluorescent intensities offluorescein-labeled pNHMMj hydrogels after 21 din PBS (with PBSexchange) shown as percentages of the initial intensities immediatelyafter gelation. Data for pNHMM2 were not available due to fastdegradation. Photos taken by Dino-Lite, intensities measured with ImageJfrom images. Scale bar=2 mm.

DETAILED DESCRIPTION

The use of numerical values in the various ranges specified in thisapplication, unless expressly indicated otherwise, are stated asapproximations as though the minimum and maximum values within thestated ranges are both preceded by the word “about”. In this manner,slight variations above and below the stated ranges can be used toachieve substantially the same results as values within the ranges.Also, unless indicated otherwise, the disclosure of these ranges isintended as a continuous range including every value between the minimumand maximum values. For definitions provided herein, those definitionsrefer to word forms, cognates and grammatical variants of those words orphrases.

As used herein, the terms “comprising,” “comprise” or “comprised,” andvariations thereof, are open ended and do not exclude the presence ofother elements not identified. In contrast, the term “consisting of” andvariations thereof is intended to be closed, and excludes additionalelements in anything but trace amounts. A “copolymer consistingessentially of” two or more monomers or residues means that thecopolymer is produced from the stated two or more monomers or containsthe stated two or more monomers and is prepared from no other monomersor contains no other residues in any quantity sufficient tosubstantially affect the LCST properties, the degradation rate in vivo,and tensile strength of the copolymer. Thus, as an example, addition ofinsignificant or trace amounts of acrylic acid or other monomers to thefeed during polymerization, or inclusion of insignificant amounts ofacrylic acid or other residues in the copolymer is considered to bewithin the scope of a copolymer consisting essentially of an N-alkylacrylamide residue in which the alkyl is one of methyl, ethyl, propyl,isopropyl and cyclopropyl; hydroxyethylmethacrylate; one or both of apolylactide-methacrylate MAPLA macromer and a HEMA-poly(trimethylenecarbonate) macromer), an N-vinylpyrrolidone monomer (VP), ahydroxyethylmethacrylate (HEMA), and/or a methacrylic acid monomer(MAA), so long as the LCST, degradation rate and tensile strength of theresultant copolymer are not significantly different than that of thesame copolymer omitting the acrylic acid residues. The significance ofeach value is determined independently and in relation to the intendeduse of the copolymer.

According to embodiments of the compounds and compositions describedherein, provided herein are injectable hydrogels that are biodegradable,elastomeric and thermoresponsive and which can easily take the shape ofa cavity into which they are injected in advance of phase transition toa solid hydrogel. The copolymers are injectable as a liquid at or belowbody temperature (about 37° C.) or room temperature (about 25° C.), orat a temperature in the range of from 10° C. to 30° C. and are solid atbody temperature. These materials are useful for a number of purposes.For example, in treatment of patients, they may be used as an injectablestem cell niche for bone marrow transplants or for other transplantationsettings; delivery vehicles for chemotherapy to tissue, such as, forexample and without limitation, gut following tumor resections; sealantsfor pulmonary and neural applications as well as for emergency treatmentof wounds. The materials also can find use as bulking agents forcosmetic applications or, even more generally, rheology modifiers. Inone embodiment, the compositions are injected in a heart for repair orregeneration of cardiac tissue.

According to certain embodiments, copolymers comprise, are preparedfrom, or consist essentially of combinations of four types ofsubunits/residues: 1) N-alkyl acrylamide in which the alkyl is methyl,ethyl, propyl, isopropyl or cyclopropyl, for exampleN-isopropylacrylamide; 2) HEMA; 3) a methacrylate-polylactide (MAPLA)macromer; and 4) a methacrylic acid (MAA). In non-limiting examples, theMAPLA macromer has a lactide:methacrylate ratio of at least 1:1, or inthe range of 2-3:1 (that is, ranging from 2:1 to 3:1). In someembodiments, the feed ratio is 75-85:5-10:3-14.5:0.5-2, wherein(NIPAAm+MAPLA):(HEMA+MAA)=85-95:5-15 (inclusive of values between thoseprovided here). In one example, the feed ratio of HEMA is 10, such thatthe feed ratio of NIPAAm:HEMA:MAPLA is 75-85:10:5-10, for example andwithout limitation, in this embodiment, the feed ratio ofNIPAAm:HEMA:MAPLA might be one of 84:10:6, 82:10:8 and 80:10:10. Inanother embodiment, the feed ratios of NIPAAm:HEMA:MAPLA:MAA are between80:5:10:5 and 80:9.5:10:0.5.

The degradation rate is positively correlated to the amount of MAAincluded in the composition. Degradation of a copolymer hydrogel formedas described herein may be 200 days and less, depending on MAA content.Those of skill will easily be able to fine-tune the MAA content to matcha preferred degradation rate. By degradation it is meant that thecopolymer (and/or hydrogel formed from said copolymer) is substantiallydegraded at the indicated time point, for example and without limitation80%, 85%, 90%, 95%, 99%, or 99.9% degraded (that is, 20%, 15%, 10%, 5%,1%, or 0.1% remaining at the indicated time point). In some embodiments,the hydrogel degrades in less than 100 days, less than 90 days, lessthan 80 days, less than 70 days, less than 60 days, less than 50 days,less than 40 days, less than 30 days, less than 20 days, less than 10days, or less than 5 days.

According to another embodiment, copolymers comprise, are prepared from,or consist essentially of four types of subunits/residues: 1) N-alkylacrylamide in which the alkyl is methyl, ethyl, propyl, isopropyl orcyclopropyl, for example N-isopropylacrylamide; 2) acrylic acid (AAc);3) a hydroxyethyl methacrylate-poly(trimethylene carbonate) (HEMAPTMC)macromer; and 4) an MAA macromer monomer. In non-limiting examples, theHEMA-poly(trimethylene carbonate) macromer has a TMC:HEMA ratio of atleast 1:1, or in the range of 2-3:1 (that is, ranging from 2:1 to 3:1).In other non-limiting examples, the feed ratio of NIPAAm:AAc:HEMAPTMC is85-87:3-5:10, for example, 86-87:3-4:10.

In other non-limiting examples, the feed ratio (the molar ratio ofmonomers in the polymerization reaction used to prepare the copolymer)of NIPAAm:HEMAPTMC is 75-85:2-14.5 (inclusive of values between thoseprovided here), with a feed ratio of MAA being in the range of 0.5-2. Inanother embodiment, the feed ratios of NIPAAm:HEMAPTMC:MAA are between80:10:5 and 80:10:0.5. The degradation rate of the copolymer is directlyproportional to the amount of MAA included in the composition. Those ofskill will understand that feed ratios of the other constituents of thecopolymer may be adjusted to any useful range. Degradation of acopolymer hydrogel formed as described herein may be 200 days and less,depending on MAA content. Those of skill will easily be able tofine-tune the MAA content to match a preferred degradation rate. In someembodiments, the hydrogel degrades in less than 100 days, less than 90days, less than 80 days, less than 70 days, less than 60 days, less than50 days, less than 40 days, less than 30 days, less than 20 days, lessthan 10 days, or less than 5 days.

In another embodiment, copolymers comprise, are prepared from, orconsist essentially of combinations of three types ofsubunits/residues: 1) N-alkyl acrylamide in which the alkyl is methyl,ethyl, propyl, isopropyl or cyclopropyl, for exampleN-isopropylacrylamide; 2) N-vinylpyrrolidone (VP); and 3) amethacrylate-polylactide (MAPLA) macromer. In non-limiting examples, theMAPLA macromer has a lactide:methacrylate ratio of at least 1:1, or inthe range of 2-4:1 (that is, ranging from 2:1 to 4:1). In someembodiments, the feed ratio for NIPAAm:VP:MAPLA is 75-85:5-20:5-10,wherein (NIPAAm+MAPLA):(VP)=85-95:5-15 (inclusive of values betweenthose provided here). In one embodiment, the feed ratio ofNIPAAm:VP:MAPLA is in the range of 70-90:5-20:5-20. In one example, thefeed ratio of VP is 10, such that the feed ratio of NIPAAm:VP:MAPLA is75-85:10:5-10, for example and without limitation, in one embodiment,the feed ratio of NIPAAm:VP:MAPLA is be one of 80:10:10 or 85:10:5. Inanother embodiment, the feed ratio of VP is 15, such that the feed ratioof NIPAAm:VP:MAPLA is 75-85:15:5-10. In one embodiment the feed ratio is80:15:5.

The degradation rate is positively correlated to the amount of VPincluded in the composition, that is, less VP leads to decreaseddegradation. Degradation of a copolymer hydrogel formed as describedherein is typically 200 days and less, depending on VP content. Those ofskill will easily be able to fine-tune the VP content to match apreferred degradation rate. As described above, by degradation it ismeant that the copolymer (and/or hydrogel formed from said copolymer) issubstantially degraded at the indicated time point, with percentagedegraded (or percentage remaining) being as described above. In someembodiments, the hydrogel degrades in less than 100 days, less than 90days, less than 80 days, less than 70 days, less than 60 days, less than50 days, less than 40 days, less than 30 days, less than 20 days, lessthan 10 days, or less than 5 days.

In addition to characterization by feed ratio, copolymers describedherein may be characterized by the ratio of incorporatedmonomer/macromer residue. For example, and without limitation,copolymers described herein may include NIPAAm:HEMA:MAPLA:MAA,NIPAAm:AAc:HEMAPTMC:MAA, or NIPAAm:VP:MAPLA in ratios similar to thosedescribed above regarding feed ratios. Those of ordinary skill in theart will understand that due to polymerization, copolymers may comprise,by molar percentage, as follows:

-   -   NIPAAm:HEMA:MAPLA:MAA 71.5-92.5:2.5-16:4.5-11:0.5-2.5, or        72-88:3-15:4.5-11:0.5-2;    -   NIPAAm:AAc:HEMAPTMC:MAA 71.5-94:2.5-16:4.5-11:0.5-2.5, or        72-88:3-15:4.5-11:0.5-2; and    -   NIPAAm:VP:MAPLA 63-93:4-22:3-22, or 68-92.5:4.5-21:3-21.        In one embodiment, the incorporated molar ratio of monomer and        macromer residues for NIPAAm:VP:MAPLA is 85-88:6-12:2-7. Those        of skill will understand that final incorporated amounts of        residues may vary from the feed ratio that is utilized by up to        10%, inclusive of values within that range, for example 3%

The copolymers, compositions and components thereof are preferablybiocompatible. By “biocompatible,” it is meant that a polymercomposition and its normal in vivo degradation products arecytocompatible and are substantially non-toxic and non-carcinogenic in apatient within useful, practical and/or acceptable tolerances. By“cytocompatible,” it is meant that the copolymers or compositions aresubstantially non-toxic to cells and typically and most desirably cansustain a population of cells and/or the polymer compositions, devices,copolymers, and degradation products thereof are not cytotoxic and/orcarcinogenic within useful, practical and/or acceptable tolerances. Forexample, a copolymer composition when placed in a human epithelial cellculture does not adversely affect the viability, growth, adhesion, andnumber of cells. In one non-limiting example, the co-polymers,compositions, and/or devices are “biocompatible” to the extent they areacceptable for use in a human or veterinary patient according toapplicable regulatory standards in a given legal jurisdiction. Inanother example the biocompatible polymer, when implanted in a patient,does not cause a substantial adverse reaction or substantial harm tocells and tissues in the body, for instance, the polymer composition ordevice does not cause necrosis or an infection resulting in harm totissues organs or the organism from the implanted compositions.

As used herein, a “polymer” is a compound formed by the covalent joiningof smaller molecules, which are referred to herein as monomers beforeincorporation into the polymer and residues, or polymer subunits, afterincorporated into a polymer. A “copolymer” is a polymer comprising twoor more different residues. Non-limiting examples of monomers, in thecontext of the copolymers described herein, include: acrylic oracrylamide monomers, acrylic N-hydroxysuccinimide ester monomers,N-hydroxysuccinimide methacrylate monomers, acrylate or methacrylateforms of N-acryloxy succinimide (NAS) monomers, hydroxyethylmethacrylate monomers, methacrylate monomers, acrylate or methacrylateforms of lactide monomers, and acrylate or methacrylate forms oftrimethylene carbonate (TMC) monomers. A monomer may be a macromerprepared from smaller monomers, such as a hydroxyethylmethacrylate-polylactide (HEMAPLA) macromer, a hydroxyethylmethacrylate-poly(trimethylene carbonate) (HEMAPTMC) macromer, amethacrylate-polylactide (MAPLA) macromer, an N-vinylpyrrolidone (VP)monomer, and/or a methacrylic acid (MAA) monomer as described herein.

Monomers (including as a group macromers) can be introduced into thecopolymer by radical polymerization or other polymerization methods,such as living polymerization (e.g., atom transfer radicalpolymerization), or in any useful manner using any suitable initiator,such as benzoyl peroxide. These polymerization processes are well-knownin the polymer chemistry field. Radical polymerization is one of themost widely used methods for preparing high polymer from a wide range ofvinyl monomers. Although radical polymerization of vinyl monomers isvery effective, it does not allow for the direct control of molecularweight, control of chain end functionalities or for the control of thechain architecture, e.g., linear vs. branched or graft polymers. Livingpolymerization systems have been developed which allow for the controlof molecular weight, end group functionality, and architecture. ATRP isa type of controlled radical polymerization or living radicalpolymerization. (see, e.g., U.S. Pat. Nos. 5,763,548, 5,807,937,5,789,487, 6,541,580, and 7,678,869). Controlled radical polymerizationmethods facilitate production of precise polymer, copolymer and blockcopolymer structures, such as A-B-A structures.

As used herein, an acrylic monomer has the general structure(CH₂═CH—C(O)—OR), and, when polymerized, forms the general polymerstructure having an alkylene backbone ( . . . C—C—C—C—C . . . ) and theoverall structure: . . . C—(—C(C(O)OR)—C—)_(n)—C(C(O)OR)—C . . . inwhich each instance of R can be the same, or in the case of a copolymer,independently different:

Polyester polymer backbones are polymer backbones containing two or moreester groups. A polyester linkage has an average of more than one esterunits (—C(O)O—), as opposed to an ester linkage that has one ester unit.An example is a methacrylate-polylactide macromer as described herein.Another example is a HEMA-poly(trimethylene carbonate) macromer. Otherexamples of residues that comprise ester linkages include, withoutlimitation, caprolactones, glycolides and a trimethylene carbonateresidues.

Polyester macromers are compounds containing on the average more thanone, and preferably two or more ester linkages. In the context ofmacromer and polymer preparations, unless otherwise indicated, thenumber of residues indicated as being present in a given polymer ormacromer is an average number and is not to be construed as an absolutenumber. Thus, as a non-limiting example, in the context of HEMAPLAmacromers, the numbers 2.1, 3.9 and 7.0 refer to an estimated averagenumber of —C(O)—CH(CH₃)—O— residues present in the macromers in themacromer composition, and, when incorporated into a copolymer, theaverage number of —C(O)—CH(CH₃)—O— residues present in the incorporatedpolyester macromer residues. The average number of residues may bedetermined by any method, for example and without limitation, by ¹H-NMR,as in the examples, below.

In describing ratios of respective monomers for any given copolymer, itis convenient to refer to feed ratios of the monomers in respect to thepolymerization method used to produce the copolymer, for example and asused herein, in reference to the radical polymerization methods used toprepare the copolymers. This is especially so when considering that theproducts of the polymerization process are polydisperse and are oftenrandom in their composition. The feed ratios typically closely representthe ratios of monomer residues in the copolymer, but typically do notexactly match because certain monomers incorporate more efficiently thanothers in any given copolymer composition. The actual ratios of monomerresidues typically vary less than 10%, and often less than 5% of thefeed ratios. As an example, in Table 1, the feed ratio of 86/4/10results in an actual composition of 88.3/3.3/8.4, a less than 3%difference in composition. As used herein a “feed ratio” refers to afeed ratio in a typical radical polymerization method, such as themethods described in the examples below and the ranges described above.

In another embodiment of the copolymer compositions described herein,the poly(NIPAAm-co-AAc-co-HEMAPTMC-co-MAA),poly(NIPAAm-co-HEMA-co-MAPLA-co-MAA), or poly(NIPAAm-co-VP-co-MAPLA)copolymers, optionally comprise an amine-reactive component or group, orare incorporated into a block copolymer with a hydrophilic polymer, suchas a polyether, which is exemplified by polyethylene glycol (PEG). Inone example, the block copolymer compositions have the structure A-B-Awhere A is poly(NIPAAm-co-AAc-co-HEMAPTMC-co-MAA),poly(NIPAAm-co-HEMA-co-MAPLA-co-MAA), or poly(NIPAAm-co-VP-co-MAPLA),and B is a polyethylene glycol block having, for example, an averagemolecular weight of from between 500 D and 25 kD, for instance between 1kD and 20 kD. The “A” blocks can be added by any useful method, forinstance, they can be synthesized by any method and attached to the Bblock by any useful chemistry. In one embodiment, the A blocks arepolymerized from the B block. The terminal portions of the B block canbe modified to act as initiators for a polymerization reaction. Asdescribed in the Examples below, the ends of a PEG block can be modifiedto act as an ATRP initiator, by addition of a suitable halide-containinggroup, for example by reacting PEG with α-bromoisobutyryl bromide. Byusing controlled radical polymerization processes, precise blockcopolymers can be prepared with low polydispersity indices (PDI), suchas PDI<2.

Lower critical solution temperature (LCST) refers to the temperaturebelow which the constituents of the hydrogel are soluble in water andabove which the constituents are insoluble. When the LCST is reached,the polymer constituents in an aqueous solution will aggregate to formhydrogel (a solid, for purposes herein). The LCST can be determined bymeasuring the change in transmittance with a UV-Vis spectrometer as afunction of temperature. LCST also can be determined by any other usefulmethod—for example and without limitation by Differential Scanningcalorimetry (DSC). DSC is used to measure LCST in the examples below.

One unique aspect of the polymers described herein is that the LCST ofthese polymers is preferably less than 37° C., and may be less than 20°C., for example, between 10° C. and about 37° C., for instance between10° C. and 25° C., so that the polymer can be distributed through themarketplace, stored and administered to a patient as a liquid at ambienttemperatures (or, if necessary, maintained at a cool temperature with anice-pack, refrigerator or other cooling device), and the polymer gels asit warms past its LCST. Many polymers suitable for administration topatients require mixing of monomers immediately prior to use, which isundesirable for many reasons. For instance, it is impractical to askdoctors, nurses or technicians to mix monomers as they need the polymer.Further, monomers can have varying degrees of toxicity. The copolymersdescribed herein do not require conducting a chemical reaction at thesite of use and the copolymers can be washed free of monomercontamination prior to distribution in the marketplace. Lastly, therelease of a portion of the aqueous phase during phase transition canfacilitate local drug delivery in the excluded aqueous phase.

Another desirable physical quality of the polymers described herein isthat, when ester linkages in the composition are hydrolyzed (forinstance over time in situ in a living system, such as a human patient),the released copolymer fragments have an LCST above 37° C., so that theyare soluble (and as an additional benefit, non-toxic), facilitating safedegradation and clearance of the polymer over time in a living systemsuch as a human body.

In one embodiment, the copolymer comprises an acrylic residue having anamine-reactive group. The copolymer may be reacted with amine-containingcompositions, such as compositions or molecules comprising amine groups,for example and without limitation, collagen, fibrin, gelatin andheparin.

In one non-limiting example in which the copolymer comprises a macromercomprising methacrylate and lactide residues, the ratio of methacrylateand lactide residues in the polyester macromer is from 1:1 to 1:10, suchas 1:1, 1:2, 1:3, 1:4, 1:5, 1:6, 1:7, 1:8, 1:9, and 1:10 (inclusive ofvalues between those provided here). In another non-limiting embodiment,the ratio of methacrylate to lactide residues in the polyester macromeris from 1:4 to 1:1, such as 1:4, 1:3, 1:2, or 1:1 (inclusive of valuesbetween those provided here). In another non-limiting example in whichthe copolymer comprises a macromer comprising hydroxyethyl methacrylateand trimethylene carbonate residues, the ratio of hydroxyethylmethacrylate to trimethylene carbonate residues in the polyestermacromer ranges from 1:1 to 1:10, 1:2 to 1:5 or any increment withinthose ranges, including 1:1, 1:2, 1:3, 1:4, 1:4.2, 1:5, 1:6, 1:7, 1:8,1:9, and 1:10 (inclusive of values between those provided here). In oneembodiment of the copolymer useful in humans or animals, the copolymerhas a lower critical solution temperature below 37° C. For veterinaryapplications, the LCST can be slightly higher as the core bodytemperature of certain animals (e.g., cats, dogs, horses, cows, sheepand goats) is in the range of 38° C.-39° C. In another embodiment, thecopolymer has a lower critical solution temperature above 37° C. afterits backbone ester linkages are hydrolyzed (substantially hydrolyzed, aswith treatment of the polymer with NaOH, as described herein).

In medical or veterinary uses, the copolymers and compositionscomprising the copolymers may serve, for example, as adhesives orfillers. They may be applied to wounds or into body cavities or used asa tissue packing to apply compression. As such, embodiments of thecopolymer solutions described herein may be applied to wounds and, inone embodiment covered, optionally with a warming compress or “heatpack” for example as are available commercially to ensure that thecopolymer is maintained at a temperature above its LCST and thus remainsgelled when in contact with any cooler areas of the body, typically theskin. As a hydrogel, embodiments of the copolymers disclosed herein maybe contained in a composition comprising the copolymer and an aqueoussolution that does not interfere substantially with the LCST and polymerstructure in its intended use. For instance, the composition maycomprise any aqueous solvent, optionally pharmaceutically acceptable,including, without limitation, water, PBS, Saline, etc. As used herein,and “aqueous solvent”, is an aqueous solution compatible with thecopolymer which can be absorbed into the copolymer matrix. Thecomposition also may comprise an active agent, biological or drug, suchas, without limitation: antibiotics, clotting agents (withoutlimitation, an antifibrinolytic, such as desmopressin/DDVAP),analgesics, anesthetics, antiseptics, anti-inflammatory agents,chemotherapeutic agents, metabolites, rheology modifiers, cytokines,chemoattractants, hormones, steroids, proteins (including enzymes),nucleic acids, cells, virus particles, nucleic acids, biomatrices orprecursors thereof, or a foaming agent. In one embodiment, thecomposition comprises stem cells (such as adipose-derived stem cells) orother progenitor cells so that the composition is useful as abiodegradable tissue engineering scaffold. The composition, even withoutcells, is useful as a cell growth niche or scaffolding into which cellssuch as native stem/progenitor cells can migrate in situ. In such anembodiment, chemokines, cellular growth agents and cellulardifferentiation agents can be included within the composition to attractcells into the composition and promote cellular growth anddifferentiation when placed in situ.

According to one embodiment, in its application to wound treatment, aclotting agent such as desmopressin may be included in a polymercomposition. An appropriate, e.g., pharmaceutically acceptable, foamingagent as are well-known in the relevant arts also may be included forthe purpose of creating compression in a wound, whether exposed to abody surface in the case of (for example) puncture wounds or bulletwounds, or internal wounds, in which case, the polymer can be injectedinto or near a site of internal bleeding. As such, the composition canfind use in many situations, ranging from home use to stabilization ofbleeding or massively bleeding patients in emergency and battlefieldsituations. The copolymer also can be used during surgical procedures toapply compression and otherwise secure a site of injury, such as aportion of a patient's intestine, nasal passage or sinus cavity where atumor or polyp has been removed or after other surgeries. The benefitsof such a reversibly-gelling copolymer composition is that thecomposition can be removed simply by cooling, for example and withoutlimitation, by flushing with cool (lower than the copolymer's LCST)flushing solution, such as water, saline or phosphate-buffered saline.Thus, while a wound and bleeding in a patient can be stabilized byapplication of the polymer, the polymer can be selectively eroded in anemergency room or during surgery simply by flushing with a cool (forexample and without limitation, 0° C. to 30° C.) saline solution.

In another embodiment, the composition as substantially described above,comprising as copolymer having a LCST of 37° C., is injected into tissueat the site of an injury or defect to provide support and/or provide ascaffold for infiltration of cells. The composition as injected mayoptionally include cells, growth factors, drugs, and the like, asprovided elsewhere in this disclosure. In certain embodiments, thecopolymer may have an LCST of, for example 36° C. or lower, 35° C. orlower, 34° C. or lower, or, in another embodiment between 10° C. and 34°C., including increments and sub-ranges therebetween, and in anotherembodiment, less than 20° C.

In certain embodiments, the composition is injected into the heart, totreat a heart defect. In some embodiments, the composition is injectedinto myocardial tissue, at the site of a myocardial defect. In someembodiments, the myocardial defect is necrotic tissue. In someembodiments, the necrotic myocardial tissue is an infarct that is theresult of a myocardial infarction. Those of skill in the art willunderstand and appreciate that the composition described above issuitable for such therapeutic uses because of its characteristics,specifically its flowable, liquid nature at room temperature (or belowthe temperature of the human body), and gel-like nature at physiologicaltemperatures (such as 37° C.). Thus, a practitioner may deliver theproper amount of the composition with precision, and without the worryof the composition “miming” into areas where it is not wanted or needed.

In a further embodiment, the composition serves as a cell growth medium.According to one embodiment, cells are introduced into a compositioncomprising a copolymer as described herein to produce a cell construct.The cell construct is incubated under conditions suitable for growth ofthe cells. That is, the cell construct can be placed in an incubator orinto a patient so that the cells are maintained under adequateenvironmental conditions to permit the cells to survive, proliferate,differentiate and/or express certain products. “Cell growth” means thatthe cells survive and preferably, though not exclusively, divide andmultiply. The composition may comprise cell growth media, whichtypically provides necessary nutrients and environmental conditions forcell growth. The cells may be introduced and incubated under conditionssuitable for cell growth by introducing the composition into a patientand allowing native cells, such as stem cells to migrate into thecomposition. The composition can be administered by injecting thecomposition into the region requiring cellular growth or remodeling,such as a region of damaged tissue.

In one non-limiting example, the damaged tissue is within the cardiacwall caused by a myocardial infarction and the composition is injectedinto the cardiac wall. In one variation of that embodiment, cytokines,chemoattractants, nutrients and/or cell differentiation factors, such asone or both of bFGF and IGF-1, are included in the composition. Thecomposition optionally contains one or more of an antiseptic, ananalgesic, an anesthetic and an antibiotic (for example, for selectionof the cells or to prevent bacterial growth in the composition). Tofacilitate cell growth, in one non-limiting embodiment, the copolymer isconjugated with collagen, for example between 0% and 10% by weight ofthe copolymer of collagen.

A current broadly pursued approach to treating ischemic cardiomyopathyis cellular transplantation into the infarct or border zone region toimprove regional and global pump function. Cells such as skeletalmyoblasts, bone marrow stromal cells, endothelial precursor cells andembryonic stem cells have been injected into injured myocardium. Thesestudies report mixed results, with modest attenuation of progressiveloss of ventricular function primarily observed in terms of maintainingor increasing LV wall thickness and fractional shortening. The mechanismbehind these beneficial results is controversial, although several havesuggested that the transplanted cells led to regeneration of contractilemyocardial tissue. Increasingly, however, it is believed that thepositive results are derived from cell-associated angiogenic effects orcytokine-mediated reduction in apoptosis rather than myocardialregeneration by the transplanted cells. In 2006, a report by Wall et al.argued that the positive results of these cell therapy studies mightsimply be attributable to the mechanical effects associated with theinjection of fluid volume (cells and delivery vehicle) into the LV wall.The injected volume would change the LV geometry and thus modify themechanics inside the LV wall, leading to a reduction of elevated localwall stresses in the infarct border zone and preventing the pathologicalremodeling in the post-infarct heart. This hypothesis was supported witha finite element analysis that modeled the local systolic fiber stressdistribution in an infarcted LV wall injected with a mechanicallypassive material. The simulation showed that injection of a volume 4.5%that of the total LV wall volume and a stiffness (elastic modulus) 20%of the natural LV tissue into the infarct border zone could decrease thefiber stress by 20% compared to a control simulation in which there wasno injection. The mechanical simulation also showed that thisattenuating effect on LV wall stress increased with the injection volumeand modulus of the injected material. This report thus provides thebasis for the local treatment of the failing cardiac wall withbiomaterial-based injection therapy. The stress reduction potential ofthe injected material is of great relevance since in a dyskinetictransmural infarct, the elevated stresses in the infarct border zoneregion are thought to contribute to pathological remodeling in thepost-infarct heart. Reducing these stresses may in turn minimizestress-induced apoptosis and border zone expansion, reducing furtherremodeling and preventing progression to congestive heart failure.

Both naturally derived and synthetic materials, including alginate,fibrin, alginate-fibrin composites, collagen, chitosan, self-assemblingpeptides, self-assembling polymers, and thermoresponsivedextran-poly(N-isopropylacrylamide) (PNIPAAm) composites, have recentlybeen utilized for cardiac wall injection therapy in animal models withreported benefits in terms of attenuated decrease in wall thickness andinfarct expansion in most cases, and in a few cases improved LVfunctions. Alginate has been shown to have a beneficial effect in termsof attenuating the decrease in wall thickness and infarct expansion, butrecent reports injecting adhesion peptide modified alginate demonstrateno clear benefit of such modification. Self-assembling peptides carryingspecific growth factors have been reported to have positive effects onthe cardiac wall remodeling process and have also been reported asvehicles for the transplantation of cardiomyocytes into the cardiacwall. Regarding thermoresponsive polymers, a recent report showed thatinjection of a dextran-poly(NIPAAm) composite 4 days following MI in arabbit model prevented adverse cardiac remodeling and dysfunction 30days following treatment.

In considering all of the biomaterials that have been utilized in theseearly investigations of cardiac injection therapy, it is encouragingthat some positive benefits have been observed in the animal modelsstudied. However, the materials investigated to date have not beenoptimal for the cardiac injection application and that mostinvestigators have utilized “off the shelf” materials (alginate, fibrin,collagen, chitosan) or synthetic hydrogels that do not display thedegradation or mechanical profile that would be most desirable for thissetting. Only short term effects have been reported in the literature,perhaps since the injected materials are rarely detectable in vivo after6 wk. Although mechanical properties of the injection material have beenshown to be important in mechanical modeling, these properties havenotably not been characterized and discussed in the early reports wherecardiac injection therapy has been investigated. In terms of the animalmodels that have been evaluated, in most reports LV injections were madewithin 1 wk of infarction, in the acute, necrotic phase. Waiting longer,even to the point of 2 wk post-MI would have greater relevance, sincethis time would more closely correspond to the beginning of the fibroticphase of remodeling, after the necrotic phase. Such a time lag maybetter represent infarcts that would be encountered in patients withsub-acute MI, where the patient may not present clinically untilsubstantial wall remodeling has already occurred.

In the example of infarcted myocardium, in addition to the mechanicalbenefits associated with injections of the copolymer compositionsdescribed herein into the infarcted myocardium, the inclusion ofbioactive growth factors in the delivered material for controlledtemporal release offers another mechanism by which injection therapymight lead to more functional LV remodeling. Many growth factors such asbasic fibroblast growth factor (bFGF), platelet derived growth factor(PDGF), hepatocyte growth factor (HGF), vascular endothelial growthfactor (VEGF) and others have been injected into the myocardiumfollowing infarction and have elicited improvements in cardiacangiogenesis, ejection fraction, and cellular activity in the form ofmitogenesis and motogenesis. Injection of fluid concentrated with growthfactors has been shown to have the same capacity to significantlyimprove cardiac function as injection of stem cells. Delivering multiplegrowth factors has also been shown to have advantages over thepresentation of a single factor. For example, cardiac injection of analginate material designed to release VEGF followed by PDGF showedincreased alpha smooth muscle cell vessel density than the delivery ofeither growth factor alone. Bimodal delivery systems may seek to mimicthe native kinetics of growth factor delivery wherein the stimulationand development of one system prior to another may be beneficial—in theexample mentioned the development first of a primary vascular networkfrom endothelial cells provided a foundation for recruiting smoothmuscle cells to mature and stabilize that network.

Two particularly important growth factors studied in the context ofcardiac remodeling have been bFGF and insulin-like growth factor-1(IGF-1). IGF-1 has been shown to have significant cardioprotective,inotropic, and regenerative capabilities and to be a potent recruitingfactor for stem cells. IGF-1 also leads to increased Akt signaling incells which can lead to production of other growth factors includingVEGF and angiopoietin-2. Local IGF-1 delivery to injured myocardium hasbeen linked to decreased apoptosis, increased cell growth, and improvedsystolic function. As such, controlled IGF-1 delivery may be useful toimprove heart function simultaneously with injected material. A growthfactor with effects complementing IGF-1 is bFGF. This potent angiogenicfactor strongly increases both endothelial and smooth muscle cellproliferation, and has been linked to increased cardiomyocyte mitoticactivity. Increased regional blood flow in the infarcted heart has beenshown as long as 6 months after a single intramyocardial injection ofbFGF. Importantly, from a functional standpoint, left ventricularejection fraction has been increased in infarcted hearts supplied withbFGF. Due to the short half-life of bFGF in vivo, controlled releasefrom biomaterial carriers has been shown to be an appropriate deliverymethod to increase cardiac regeneration. Using a bimodal delivery systemof bFGF followed by IGF-1 may provide a vascular network to which stemcells can be recruited followed by increased proliferation with animproved local vascular network.

In addition to the above description of cardiac uses of the copolymersdescribed herein, there are numerous other uses for supplementing and/orenhancing repair of other muscle tissues. Those of ordinary skill in theart will appreciate that the presently disclosed copolymers will besuitable for numerous applications where biocompatible gels may beuseful.

Compositions comprising a copolymer described herein can be distributedfor use in any suitable vessel. In one instance, the composition ispackaged in a sealed container, from which the composition can bepoured, squeezed or otherwise decanted, for example and withoutlimitation, by use of a syringe. The vessel can be a bag, such as an IVbag. In another embodiment, the composition can be distributed in asyringe for immediate dispensation into a wound or body cavity/location.A syringe can be fitted with any type of needle, tip, tube, balloondevice or other useful fitting for facilitating accurate placement ofthe solution in or around a desired delivery site, for example andwithout limitation, for delivery into the large intestine of a patientafter removal of a tumor. In another embodiment, the composition and apharmaceutically acceptable solvent is stored within a syringe at orbelow 4° C. and the syringe is fitted with a needle gauge sufficient toallow for injection without increased pressure but also prohibit backflow of the solution into the syringe after injection, such as, withoutlimitation, a 16 G through 23 G (gauge) needle, and in certainembodiments an 18 G or 20 G needle. As described below and in theExamples, a robotic injection device can be used to deliver any of thecompositions described herein to the heart or other organs or tissue.Thus, methods of use embodying the above-described uses for a copolymerdescribed herein and compositions comprising the copolymer arecontemplated and embraced as part of the present invention.

In the context of myocardial infarction, although myocardial injectiontherapy is currently dominated by transcatheter endocardial approaches,direct epicardial injection offers potential advantages such as easydetection of target myocardial infarct lesions, decreased likelihood ofcerebrovascular complications, and superior site specific efficacy.Particularly with gel materials, the risk of backflow and embolizationfrom an endocardial injection site is a serious concern. To date, amajor limitation of direct epicardial injection is the lack of dedicatedminimally invasive access technology, generally causing it to beperformed only in conjunction with other procedures using sternotomy orthoracotomy, both of which have high associated morbidity. In addition,the instrumentation used in most reported applications does not readilyaccommodate the motion of the beating heart, and therefore does notfacilitate precise placement and depth of injections. A dedicatedtechnology for precise interaction with the heart from within theintrapericardial space that balances treatment efficacy and minimalinvasiveness is likely to provide a future clinical benefit for thehydrogel injection therapy proposed here and for myocardialinjection-based therapies in general. To address this need, we havedeveloped a novel miniature robotic device (HeartLander, see, e.g., USPatent Publication No. 20050154376, incorporated herein by reference inits entirety) that navigates over the epicardial surface to performminimally invasive myocardial injections on the beating heart through asubxiphoid approach. Such injections have been achieved in vivo in aporcine model, demonstrating positioning accuracy of 1.7±10 mm inapplying multi-target injection patterns.

In another use, a composition described herein can be used for cosmeticpurposes, such as for a rheology modifier. Ingredients, includingwithout limitation colorants, fragrances, flavors, and other ingredientslisted herein, including active agents, may be included in thecomposition.

The following examples are provided for illustration purposes and arenot intended to limit the scope of the present invention. Reference ismade to International Patent Publications: WO 2008/045904 and WO2010/127254, for the disclosure of the general process of making, andcharacteristics of, various copolymers. The disclosures of WO2008/045904 and WO 2010/127254 are incorporated herein by reference intheir entirety.

EXAMPLES

A hydrogel possessing thermoresponsive behavior coupled with robustmechanical properties suitable for soft tissue engineering is of greatinterest. Such a thermoresponsive scaffold could readily encapsulate anddeliver cells for subsequent mechanical training in vivo or in vitro.Described herein and in the examples below is a family of injectable andflexible hydrogel composites based on thermosensitive copolymers,optionally conjugated with collagen. The compositions find use in, forexample cardiac remodeling after myocardial infarction. These novelthermosensitive, biodegradable and flexible hydrogels have propertiesattractive for future application in soft tissue engineering.

Example 1 A Thermally Responsive Injectable Hydrogel Incorporatingacrylic acid-poly(trimethylene carbonate) for Hydrolytic Lability

Thermally responsive injectable and bioabsorbable hydrogel bycopolymerization of N-isopropylacrylamide (NIPAAm), acrylic acid (AAc),and biodegradable monomer hydroxyethyl methacrylate-poly(trimethylenecarbonate) (HEMAPTMC) is synthesized and evaluated. The study sought toinvestigate and tune the molecular design by altering the relativeamount of AAc so that a thermoresponsive hydrogel would be achieved withan LCST below body temperature prior to hydrolysis of thepoly(trimethylene carbonate) (PTMC) branches, but with an LCST that roseabove body temperature with PTMC cleavage. The HEMAPTMC component wasselected and synthesized for use since the carbonate bond in PTMC shouldhave a hydrolysis rate that would allow retention of the gel over theseveral week period that are hypothesized as being necessary for thecardiac application in vivo. After characterizing and optimizing thecopolymer structure, the optimized hydrogel was evaluated by injectioninto chronic rat myocardial infarctions two weeks following coronaryligation, and the resulting cardiac performance and ventricularremodeling were assessed over an 8 week period. The hypothesis was thatinjection of the designed thermoreponsive hydrogel would alter theprogression of ventricular remodeling, preserving ventricular wallthickness and maintaining contractile function.

Chemicals were purchased from Sigma-Aldrich unless otherwise stated.NIPAAm was purified by recrystallization from hexane and vacuum dried.NIPPAm (50 g) was dissolved into 150 mL hexane at 80° C. and thenrecrystallized at room temperature. AAc and 2-hydroxyethyl methacrylate(HEMA) were purified by vacuum distillation at 70° C. and 100° C.,respectively. Benzoyl peroxide (BPO), stannous 2-ethylhexanoate[Sn(OCt)₂], trimethylene carbonate (TMC, Boehringer Ingelheim ChemicalsInc.) were used as received.

Synthesis of HEMA-polyTMC (HEMAPTMC)

HEMAPTMC was synthesized by ring-opening polymerization of TMC initiatedby HEMA with Sn(OCt)₂ as a catalyst (FIG. 1). Stoichiometric amounts ofHEMA and TMC (molar ratio 1:2) were mixed in a flask to which was addedanhydrous toluene of equal mass to the TMC/HEMA mixture. Sn(OCt)₂ (1 mol% with respect to HEMA) in 1 mL toluene was subsequently added. Thereaction was conducted at 120° C. for 1.5 h. The mixture was thendissolved in THF and precipitated in water. This precipitation processwas repeated twice and the liquid precipitate was then isolated bycentrifugation, dissolved in THF, and dried over anhydrous MgSO4. THFwas removed by rotary evaporation.

Synthesis of poly(NIPAAm-co-AAc-co-HEMAPTMC)

Poly(NIPAAm-co-AAc-co-HEMAPTMC) copolymers were synthesized by freeradical polymerization (FIG. 1). Monomers (NIPAAm, AAc, HEMAPTMC) weredissolved in 1,4-dioxane to form a 5 wt % solution containing BPO(7.2×10⁻³ mol/mol monomer). The polymerization was carried out at 70° C.for 24 h under argon atmosphere. The copolymer was precipitated inhexane and further purified by precipitation from THF into diethylether. The purified copolymer was vacuum dried.

Results Synthesis of HEMAPTMC and Copolymer

The synthesis of HEMAPTMC was confirmed by the 1H-NMR spectrum of theproduct (data not shown) and the ¹³C-NMR spectrum (data not shown)containing proton peaks and carbon peaks in agreement with the molecularstructure of HEMAPTMC. In the ¹H-NMR spectrum, HEMA alone would beexpected to have two characteristic triple peaks centered at 4.4 ppm and3.9 ppm for d protons, while for HEMAPTMC the combination of the two dpeaks into a single peak at 4.4 ppm provides confirmation of theformation of HEMAPTMC. The chemical structure of HEMAPTMC was furtherconfirmed by the mass spectrum (API-ES positive). Peaks at 254.8(HEMAPTMC1+Na⁺), 357.0 (HEMAPTMC2+Na⁺), 459.0 (HEMAPTMC3+Na⁺), 561.0(HEMAPTMC4+Na⁺) and 663.0 (HEMAPTMC5+Na⁺) were observed, indicating thatthe product was a mixture of molecules containing different PTMC lengths(data not shown). The number average length of PTMC units per monomerwas determined from ¹H-NMR spectrum as 2 by calculation from the ratioof the integrals of hydrogen peaks from PTMC and the double bondhydrogen (CH2=) peak. This PTMC unit number for HEMAPTMC was inagreement with the molar feed ratio of HEMA to TMC (1:2) in thesynthesis of HEMAPTMC.

Copolymers with different monomer ratios were prepared by free radicalpolymerization (FIG. 1). Table 1 (below) summarizespoly(NIPAAm-co-AAc-co-HEMAPTMC) copolymers synthesized with differentAAc feed ratios. All of the copolymers have molecular weights between 20k and 30 k, and a polydispersity index of 1.5˜2.0. The existence of AAc(—COOH) units in the copolymer was verified and quantified by titrationof the polymer solution with NaOH solution (0.1 M) (data not shown). TheAAc content obtained by the titration method and the integration ratiosof characteristic proton peaks in the ¹H-NMR spectra were used todetermine copolymer compositions (Table 1 below). The monomercompositions in the copolymers were found to be close to the feedratios, with a consistent slight reduction in the measured AAc contentfrom that expected based on the feed ratio.

TABLE 1 Properties of poly(NIPAAm-co-AAc-co-HEMAPTMC) copolymers withdifferent feed ratios of AAc. LCST —COOH Polymer 37° C., 16.7 wt 16.7 wt% Feed ratio Mn Mw/ content, composition, % in PBS, in PBS,NIPAAm/AAc/HEMAPTMC Yield g/mol Mn 10⁻⁴ mol/g NIPAAm/AAc/HEMAPTMC pH 7pH 7, ° C. 87/3/10 86% 27,000 1.8 1.6 88.5/2.1/9.4 solid gel 29.1 ±0.37* 86/4/10 87% 23,000 1.9 2.6 88.3/3.3/8.4 solid gel 33.1 ± 0.43*85/5/10 84% 34,000 1.5 2.8 87.0/3.6/9.4 cloudy, weak 36.2 ± 0.38* gel84/6/10 93% 21,000 2.0 3.8 86.2/4.8/8.9 clear solution 44.5 ± 0.10* p <0.001 versus each of other copolymers

Gelation Properties, the LCST and Optimization of Monomer Feed Ratio

The qualitative gelation properties of thepoly(NIPAAm-co-AAc-co-HEMAPTMC) copolymers are summarized in Table 1(above). When the AAc feed ratio was 3% and 4%, a solid gel could beformed at 37° C. When the AAc feed ratio was increased to 5%, afluid-like hydrogel with negligible strength was formed. When the AAcfeed ratio was as high as 6%, the copolymer solution remained a clearsolution at 37° C., indicating an LCST above 37° C. The calculated LCSTswere determined from the optical data (not shown). The temperature atwhich optical absorption rapidly transitions (the LCST) is seen toincrease as the AAc feed ratio of the copolymer is increased. Whilecopolymers with AAc feed ratios of 3,4 and 5% had LCSTs below 37° C.,the copolymer with an AAc feed ratio of 6% had an LCST of 45° C.

Example 2 A Thermally Responsive Injectable Hydrogel IncorporatingMethacrylate-Polylactide for Hydrolytic Lability

Methacrylate-polylactide (MAPLA), with an average 2.8 lactic acid units,was synthesized and copolymerized with n-isopropylacrylamide (NIPAAm)and 2-hydroxyethyl methacrylate (HEMA) to obtain bioabsorbable thermallyresponsive hydrogels. Poly(NIPAAm-co-HEMA-co-MAPLA) with three monomerfeed ratios (84/10/6, 82/10/8 and 80/10/10) was synthesized andcharacterized with NMR, FTIR and GPC. The copolymers were soluble insaline at reduced temperature (<10° C.), forming clear solutions thatincreased in viscosity with the MAPLA feed ratio. The copolymersunderwent sol-gel transition at lower critical solution temperatures of12.4, 14.0 and 16.2° C. respectively and solidified immediately uponbeing placed in a 37° C. water bath. The warmed hydrogels graduallyexcluded water to reach final water contents of ˜45%. The hydrogels asformed were mechanically strong, with tensile strengths as high as 100kPa and shear moduli of 60 kPa. All three hydrogels were completelydegraded (solubilized) in PBS over a 6-8 month period at 37° C., with ahigher MAPLA feed ratio resulting in a faster degradation period.Culture of primary vascular smooth muscle cells with degradationsolutions demonstrated a lack of cytotoxicity. The synthesized hydrogelsprovide new options for biomaterial injection therapy where increasedmechanical strength and relatively slow resorption rates would beattractive.

Materials and Methods

All chemicals were purchased from Sigma-Aldrich unless otherwise stated.NIPAAm was purified by recrystallization from hexane and vacuum dried.2-hydroxyethyl methacrylate (HEMA) was purified by vacuum distillation.Lactide was purified by recrystallization from ethyl acetate. Benzoylperoxide (BPO), sodium methoxide (NaOCH₃) and methacryloyl chloride wereused as received.

Synthesis of Methacrylate Polylactide (MAPLA)

As shown in FIG. 2, polylactide (HO—PLA-OCH₃) was synthesized by NaOCH₃initiated ring opening polymerization of lactide. In a lactide solutionin dichloromethane a solution of NaOCH₃ in methanol (10% wt/v) was addedwith a molar ratio of (NaOCH₃+HOCH₃) to lactide of 1:1, under vigorousstirring. The reaction proceeded for 2 h at 0° C. before the solutionwas rinsed with 0.1M HCl and deionized (DI) water. The organic phase wasisolated by centrifugation and dried over anhydrous MgSO₄. The solvent(dichloromethane) was removed by rotary evaporation at 60° C. to obtainHO—PLA-OCH₃. Biodegradable monomer MAPLA was synthesized by droppingequimolar amounts of methacryloyl chloride into the HO—PLA-OCH₃ solutionin dichloromethane in the presence of equimolar amounts of triethylamineAfter reacting at 0° C. overnight, the solution was filtered to removeprecipitants, and was then rinsed sequentially with 0.2M Na₂CO₃, 0.1MHCl and DI water. The organic phase was isolated by centrifugation anddried over anhydrous MgSO₄. The solvent (dichloromethane) was removed byrotary evaporation at 40° C. to get the raw product of MAPLA, which wasfinally purified by flash chromatography.

Synthesis of poly(NIPAAm-co-HEMA-co-MAPLA)

Poly(NIPAAm-co-HEMA-co-MAPLA) copolymers were synthesized by freeradical polymerization (FIG. 2). Monomers (NIPAAm, HEMA, MAPLA) weredissolved in 1,4-dioxane to form a 5 wt % solution containing BPO(7.2×10⁻³ mol/mol monomer). The polymerization was carried out at 70° C.for 24 h under argon atmosphere. The copolymer was precipitated inhexane and further purified by precipitation from THF into diethyl etherand vacuum dried.

Example 3 Design Rationale and Characterization ofpoly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA)+/−Growth Factors

In order to provide the capacity for controlled release of bioactivefactors from the injected hydrogels, the design ofpoly(NIPAAm-co-HEMA-co-MAPLA was modified to createpoly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA by incorporating the monomerN-hydroxysuccinimide methacrylate (MANHS) (FIG. 3). This monomer has theability to readily react with primary amine groups (e.g. surface lysinesin proteins) forming a stable amide bond and provides a means forcovalent growth factor attachment to the copolymer in an aqueousenvironment. The covalent link reduces the burst release oftenencountered with hydrogel systems. This attachment technique was appliedpreviously, where poly(NIPAAm-co-AAc-co-HEMAPTMC) was used but with theaddition of N-acryloxy succinimide (NAS) to bind IGF-1. The resultsshowed IGF-1 was successfully bound to the hydrogel and remainedbioactive upon release. As discussed, however,poly(NIPAAm-co-AAc-co-HEMAPTMC) does not have the mechanical strengthand decreased degradation rate advantages ofpoly(NIPAAm-co-HEMA-co-MAPLA). MANHS may be desirable over NAS for thisapplication since the succinimide ester reactivity with water is slowerfor MANHS, ultimately favoring amine reaction and higher loadingefficiency. In preliminary studies 1 mol % MANHS has been incorporatedin the polymer feed to make poly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA) (Mw25 kD, Mw/Mn ˜1.5) that was subsequently loaded with protein at aloading efficiency of 46%. The majority of protein was delivered invitro from poly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA) in the first weekfollowed by near zero-order release extending for 3 months related topolymer degradation. Because we ultimately seek a hydrogel system withbi-modal release, with bFGF delivery occurring before IGF-1, we willtake advantage of the higher early release rates and use this covalentattachment system with bFGF.

Several design parameters of poly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA) canbe altered to influence protein release kinetics. In one example, therelative amount of MANHS incorporated into the copolymer, whichdetermines protein binding capacity, is varied between 1-5 mol %.Increasing MANHS content not only has the potential to increase proteinbinding but can also be used to speed polymer degradation and thusprotein delivery rate. In addition to this parameter the amount of bFGFloaded can be varied. Both parameters will influence the release profileof the protein. While studies to date have not specifically investigatedthe influence that burst release has on the angiogenic effects ofdelivered bFGF, some conclusions can be made based on direct injectionstudies. It has been shown that when a solution of free bFGF is injecteddirectly into the myocardium only 16% remains after one hour. While somecardiac improvements have been shown from this delivery method,additional benefits of bFGF delivery with a carrier have beendemonstrated. Delivery of bFGF from microspheres and gel systems over arange of 1-6 wk resulted in substantial vascular and functionalimprovements. The amount of bFGF remaining at the injection site 72 hafter injection was roughly 30% when the factor was incorporated intogelatin microspheres —a 15× increase compared to free bFGF injection. Asan example, the design of poly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA) ismanipulated with the objective of >70% release over the first 2 wk, withcontinued delivery for at least 6 wk.

Studies have also shown that improvements to cardiac function and bloodflow in rats can be elicited when between 10 and 100 μg of bFGF isdelivered, providing an exemplary range of bFGF loading concentrationsthat can be characterized in vitro. While there is concern thatexcessive bFGF delivery might lead to hemangioma formation, bFGFadministration in this range has not been associated with thiscomplication in animal studies.

In another example microparticle carriers, which offer an extendedrelease profile, are employed as a protein delivery mechanism. Combininga suspension of growth factor-loaded (or other active agent-loaded)microparticles in solution with the protein-conjugatedpoly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA) permits delivery of a secondgrowth factor, where each delivery system—covalent hydrogel attachmentor microparticulate —has distinct design parameters to influence releasekinetics which are largely independent of the other. Microparticles ofmany common biomaterials such as gelatin, collagen, alginate, andpoly(lactic-co-glycolic) acid (PLGA) have been synthesized and used fordrug delivery with positive results. Withpoly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA). In an example, relativelyhydrophobic PLGA microparticles are utilized since they will interactwith the hydrophobic NIPAAm groups in the collapsed hydrogel, thusprecluding their exclusion from the gel network during phase transition.As an example, a double emulsion system is used to form thesemicroparticles with IGF-1 loading in a manner to protect against proteindenaturation, as previously described. It has been shown thatappropriately designed microparticles can deliver growth factor at aslower rate with a smaller burst than the covalent system above. PLGA(75:25, 100 kDa) microparticles (49 um diam) encapsulating BSA have beensynthesized and protein release rates measured after particle mixingwith poly(NIPAAm-co-HEMA-co-MAPLA). Inclusion of protein-loaded PLGAmicroparticles in a hydrogel system has previously been shown to nearlyeliminate burst release leading to delayed protein delivery. The resultsagree, showing only 4% burst release of total protein during gelformation, and only 15% released by 2 wk. This release rate is aboutone-third of that from the same microparticles not within a hydrogel.Later-stage protein release follows the degradation of the PLGAmicroparticles which increases after 4 wk in saline. A biphasic systemis thus achievable wherein the majority of bFGF is released early fromthe hydrogel carrier followed by IGF-1 release later as the PLGAmicroparticles within the gel degrade. As has been shown previously, theburst and duration of protein release from PLGA microparticles can fallwithin a wide range depending on controllable factors such as polymerweight fraction in the microparticles, particle size, degradation time,and weight fraction of growth factor. One exemplary design objective formicroparticles is <20% release in the first 2 wk, with an additional 60%over the 4 wk following. Since a dose range from 25 ng to 100 μg ofIGF-1 has shown functional cardiac improvements in rats, a moderatetotal dose between 1-10 μg is used.

Hydrogel chemical structure is characterized with NMR, FTIR, and massspectra. Molecular weight is determined by gel permeationchromatography. The LCST of the hydrogel solutions is determined by DSC,UV-optical absorption with temperature scanning and rheological testingwith temperature scanning. Hydrogel solution viscosity below the LCST ismeasured with rheometry and gelation speed at 37° C. is quantified byplotting water content over time. Tensile and rheological testingprovides hydrogel mechanical properties. Polymer degradation productcytotoxicity is assessed by the metabolic viability of cells culturedwith medium supplemented with degradation products. Cells also areobserved under fluorescence microscopy after live/dead staining whencultured atop the hydrogels. For controlled release frompoly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA), the attachment of bFGF to thepolymer is investigated with matrix assisted laser desorption ionization(MALDI) mass spectrometry. Release kinetics of each growth factor fromits polymer carrier is analyzed by enzyme-linked immunosorbant assay forthe specified protein. To quantify bioactivity of the released bFGF andIGF-1 cell proliferation assays with L929 fibroblasts and MG-63 cellsare used, respectively with calibration to known growth factorconcentrations. Failure to meet the stated design objectives forpoly(NIPAAm-co-HEMA-co-MAPLA) and poly(NIPAAm-co-HEMA-co-MANHS-co-MAPLA)results in iterative material design refinement and characterizationusing the controllable parameters discussed above.

Example 4 Tailoring the Degradation Rates of Thermally ResponsiveHydrogels Designed for Soft Tissue Injection by Varying theAutocatalytic Potential

Appropriate biomaterial degradation behavior is essential for obtainingdesired therapeutic outcomes in a variety of tissue engineering andregenerative medicine applications. Biomaterial degradationtheoretically should be aligned with the pace of cell infiltration andneo-tissue formation to allow the structural and functional integrationof host tissue with tissue developed in the region of the implantedbiomaterial. For example, rapid degradation of dermal grafts may favor afibrotic response over a more constructive regeneration outcome.Uncoordinated absorption of bone substitutes can cause mechanicalmismatch and ultimately contribute to failure in load bearing. Inbiodegradable arterial stent development, there is interest in assuringthat the stent remains long enough to remodel the vascular wall in astable fashion, but not much longer to avoid complications associatedwith a permanent foreign body in the vascular wall.

Thermally responsive hydrogels have been widely studied for theiramenability to minimally invasive delivery in the realms of drugdelivery, embolization therapy, cell delivery vehicles, tissue fillersand wound dressings. More recently, intramyocardial injection therapyhas been pursued using mechanically strong thermally responsivehydrogels to inhibit pathological ventricular dilatation aftermyocardial infarction, a major contributor to morbidity and mortality inischemic cardiomyopathy. As with other biomaterial applications wheretemporary mechanical support is the objective, the degradation behaviorof thermally responsive hydrogels becomes a critical designconsideration.

In seeking to control the degradation rate for a thermally responsivehydrogel that would be used in soft tissue injection, several otherdesign criteria must be considered. Basic requirements are acceptablylow levels of cytotoxicity of the polymer and degradation products andadequate thermal response to allow needle-based injection and stiffeningin situ. In the role of cell carrier, one would want to maintainencapsulated cell viability. When tuning the degradation rate, one wouldideally not impair the thermal sensitivity of the system orsubstantially alter the mechanical properties in the fluid or hydrogelstate. Polyesters, widely used as biodegradable polymers in general,have been utilized as a hydrophobic component in thermally responsivehydrogels to trigger dissolution and absorption of the hydrogels uponester cleavage. Polyester materials degrade faster under low pHconditions due to catalyzed hydrolysis. The accumulation of acidicdegradation products can lead to an autocatalytic effect, acceleratingthe hydrolysis of ester bonds. It was anticipated that the autocatalysiseffect could be employed to tune degradation rates across wide ranges.Thus, in this study the amount of acid in a thermally responsivecopolymer backbone was varied to modulate the degradation rate of apoly(N-isopropylacrylamide) based hydrogel,poly(NIPAAm-co-HEMA-co-MAPLA) (pNHM, copolymerized withN-isopropylacrylamide (NIPAAm), 2-hydroxyethyl methacrylate (HEMA) andmethacrylate-polylactide (MAPLA)). The pendant hydrophobic MAPLAsidechains become acidic units upon hydrolysis, resulting in a highertransition temperature and eventual solubility of the copolymer, withoutbackbone cleavage. Different molar ratios of methacrylic acid (MAA) wereincorporated into the copolymer to obtainpoly(NIPAAm-co-HEMA-co-MAPLA-co-MAA) (pNHMMj). The effect of MAA onhydrogel degradation was studied as well as its effect on thermal andmechanical behavior. The cytotoxicity of pNHMMj hydrogels and theirdegradation products were evaluated and an in vivo degradation study ofpNHMMj hydrogels was performed in a rat hindlimb injection model.

Materials and Methods

All chemicals were purchased from Sigma-Aldrich unless otherwise stated.N-isopropylacrylamide (NIPAAm) was purified by recrystallization fromhexane and vacuum-dried. 2-Hydroxyethyl methacrylate (HEMA) was purifiedby vacuum distillation. Lactide, benzoyl peroxide (BPO), sodiummethoxide (NaOCH₃), methacryloyl chloride, methacrylic acid (MAA) andother solvents were used as received.

The synthesis of methacrylate polylactide was performed as previouslydescribed (Ma et al. Thermally responsive injectable hydrogelincorporating methacrylate-polylactide for hydrolytic lability.Biomacromolecules 2010; 11: 1873-81). Briefly, NaOCH₃/methanol was addedto a lactide/dichloromethane solution to synthesize polylactide(HO—PLA-OCH₃) through ring-opening polymerization. MAPLA was synthesizedby dropping methacryloyl chloride into a HO—PLAOCH3/dichloromethanesolution containing triethylamine Dichloromethane was removed by rotaryevaporation and the product was purified by flash chromatography toobtain MAPLA with yields of —60%.

Poly(NIPAAm-co-HEMA-co-MAPLA-co-MAA) (pNHMMj) copolymers (FIGS. 4 and 5)were synthesized from NIPAAm, HEMA, MAPLA and MAA by free radicalpolymerization. The feed ratios of NIPAAm, HEMA, MAPLA and MAA were80/(10-j)/10/j, where j=0, 0.5, 1, 2, 5, 10 (FIG. 4). Table 2, showingthe parameters for production illustrated in FIG. 4, is shown below.

TABLE 2 Feed ratio for various poly(NIPAAm-co- HEMA-co-MAPLA-co-MAA)(pNHMMj) copolymers MAA feed Feed ratio (%) ratio (%) NIPAAm HEMA MAPLAMAA 0 80 10 10 0 0.5 80 9.5 10 0.5 1 80 9 10 1 2 80 8 10 2 5 80 5 10 510 80 0 10 10

Monomers (0.066 mol) were dissolved in 180 mL of 1,4-dioxane containing0.23 g BPO. The polymerization was carried out at 75° C. for 20 h underargon protection. The copolymer was precipitated in hexane and furtherpurified by precipitation from THF into diethyl ether and vacuum-dried,with yields of ˜80%. Fluorescently labeled copolymers were synthesizedusing the same reaction conditions with fluorescein O-methacrylate addedat a feed ratio of an additional 2%, with all of the other monomer molarratios remaining constant and j=0. Fluorescently labeled hydrogels usedin the in vivo study were prepared by dissolving 14.25 wt % unlabeledcopolymer with 0.75 wt % labeled copolymer in PBS.

¹H NMR spectra of pNHMMj were recorded with a 600 MHz Brukerspectrometer using CD3Cl or DMSO-d6 as a solvent. Molecular weight ofthe copolymers was determined by gel permeation chromatography (GPC,Waters Breeze System, Waters 1515 HPLC Pump, Waters 2414 differentialrefractometer). The copolymers were dissolved in THF at a concentrationof 1 mg/mL and the GPC analysis was performed at 35° C. A poly(methylmethacrylate) standard kit (Fluka, ReadyCal Set Mp 500-2700000) was usedfor molecular weight-elution volume calibration.

Rheology studies were conducted on a TA Instruments rheometer (AR2000)to observe viscosity changes in the hydrogels during the temperatureinduced solegel transition. The polymer solutions (15 wt % in PBS) wereplaced between two parallel plates. With a temperature sweep from 5 to35° C. and a heating rate of 5° C./min, the shear storage modulus G0 andthe loss modulus G00 were collected as a function of temperature at afixed strain of 2% and a frequency of 1 Hz.

To measure the mechanical properties of the hydrogels, samples wereincubated in a 37° C. water bath for 24 h to reach a stable watercontent, and then the solid hydrogels were cut into rectangular strips 1mm thick, 4 mm wide, and 25 mm long and then loaded in a water bathequilibrated to 37° C. An ElectroForce 3200 Series II (Bose, Minnesota,US) equipped with a 2.5 N load cell was utilized to record the tensilestress-strain curve immediately after the samples were taken out of thewater bath.

Hydrogel degradation was quantified by mass loss measurements. Hydrogelswith known initial dry masses (—60 mg) were immersed into 6 mL of PBS at37° C. At predefined time points over a 28 week period the hydrogels(n=3 each) were lyophilized and the relative mass loss was recorded. ThepH of the supernatant during degradation was measured with an Accumet pHmeter (Fisher Scientific, Waltham, Mass.). To compare the relative pHinside pNHMMj hydrogels, polymers were dissolved in PBS containing 2.5mM LysoSensor Yellow/Blue DND-160 pH sensitive dye (Life Technologies,Grand Island, N.Y., US). After gelation, the excess fluid was removedand replaced with PBS. The hydrogels were allowed to stabilize for 1 dbefore being placed in a plate reader with excitation at 360 nm andemission intensities at 440 nm and 540 nm measured for the hydrogelsurface and cross sections (cut and exposed). The cross sections wereused to semi-quantitatively determine whether there was detectably lowerpH in the hydrogel interior.

The cytotoxicity of the pNHMMj degradation products was assessed bymeasuring the relative metabolic activity of rat vascular smooth musclecells (rSMCs) cultured in Dulbecco's modified Eagle medium (DMEM)(Gibco, Life technologies) with 10% fetal bovine serum (FBS), 1%penicillin/streptomycin, and supplemented at 10% with hydrogeldegradation solution. The hydrogel degradation solution was preparedfrom incubation of the hydrogel in PBS. Culture medium with PBS added at10% was used as a negative control rSMCs were seeded at an initialdensity of 30,000/cm2 and their metabolic activity was measured (n=each)using an MTS assay kit (Promega CellTiter 96 Cell Proliferation Assay).To qualitatively verify the results of the above test, cells were alsoobserved under fluorescence microscopy after live/dead staining with aPromokine Live/Dead Cell Staining Kit.

rSMCs labeled with the live cell marker CellTracker Red CMTPX (LifeTechnologies, Grand Island, N.Y.) were suspended in PBS at a density of2×107/mL. A total of 0.25 mL of this cell suspension was then added into1 mL of the hydrogel solutions at 4° C. The mixture was thoroughly mixedbefore being transferred into a 37° C. water bath for gelation. Thesupernatant was removed and replaced with culture medium (DMEMsupplemented with 10% FBS, and 1% penicillin/streptomycin). The mediumwas changed every 3 d. After 1 and 7 d of culture, samples were takenout and cut into 100 mm thick sections and observed directly underfluorescence microscopy (Eclipse Ti-U, Nikon Instruments). At 7 danother set of samples were cooled at 4° C. to release the encapsulatedcells and these released rSMCs were then cultured on TCPS for anadditional 7 d to qualitatively assess their ability to proliferate. Ina separate set of experiments, unstained rSMCs were encapsulated asdescribed above and upon recovery from the hydrogels by cooling at 1 and4 d were stained with trypan blue solution and the percentages of livecells were calculated by manual counting of multiple microscopic fieldsfor 3 independent samples for each hydrogel type.

Adult female Lewis rats weighing 160-210 g were utilized in a protocolthat followed the National Institutes of Health guidelines for animalcare and that was approved by the University of Pittsburgh'sInstitutional Animal Care and Use Committee. Anesthesia was induced with3.0% isoflurane inhalation with 100% oxygen followed by 1.5e2%isoflurane with 100% oxygen during procedure. Dermatotomy was performedto expose the inner thigh muscles on both legs. Single injections of200-250 mL of hydrogel (fluorescently labeled or unlabeled) were madeapproximately 3 mm deep in the muscle bed. For each hydrogel, 6injections in 6 legs were made (4 labeled, 2 unlabeled). Inner thighmuscles from 2 legs of the labeled groups were excised 3 min afterinjection. The muscles were incised to expose the hydrogels, and imageswere taken with a Dino-Lite (AM4113T-GFBW, AnMo Electronics, New TaipeiCity, Taiwan) under bright field and fluorescence mode. After 21 d, ratswere sacrificed and the inner thigh muscles encompassing the hydrogelswere excised, images taken and the tissue was fixed in 10% formaldehydefor 3 d before embedding. H&E staining (for unlabeled hydrogels) andimmunohistochemical staining (with labeled hydrogels) with monoclonalantibodies against CD68 (1:100, Abcam) was performed. Nuclei werestained with 400,6-diamidino-2-phenylindole (DAPI; 1:10000, Sigma).Microscopic images were taken under fluorescence microscopy and assessedwith ImageJ.

For the paired comparisons of FIG. 7, a paired t-test was employed.Where three or more groups were being compared, one-way ANOVA wasemployed with Tukey's test applied for specific comparisons. Results arepresented as the mean with standard deviation. Statistical significancewas defined as p<0.05.

Results

The incorporation of MAA into the hydrogels was confirmed by NMR as theeCOOH peak appeared at ˜12 ppm on the ¹H spectrum, as shown in FIG. 6A.The content of MAA in the copolymer calculated with peak areas shows alinear increase with the MAA feed ratio, (FIG. 6B) although this contentis ˜40% lower than the MAA feed ratio in the reaction system. Thisresult was also confirmed with minor discrepancies from the NMR resultsby titrating the eCOOH groups in the copolymer with HCl/NaOH forprotonation/deprotonation (FIG. 6B). GPC results showed that themolecular weights Mw of the pNHMMj copolymers were all between 22000 and26000 g/mol.

The pH of the supernatant solution after gelation of the pNHMMj polymersin PBS buffer (15 wt % of copolymers) showed that with increased MAAcontent in the copolymer, the more acidic the initial degradationenvironment was for PLA sidechains in this limited volume system, FIG.7. Measured immediately after changing PBS, the supernatant pH for allpNHMMj increased, and showed no significant difference compared to thepH of PBS. After stabilization for 24 h to allow diffusion, thesupernatant pH of pNHMM0, pNHMM0.5 and pNHMM1 remained above 7.3,whereas the pH for pNHMM2 dropped below 7.1 (FIG. 25, pNHMM5 and pNHM10degraded too quickly for this measurement). After another cycle of PBSchange and measurement after another 4 d, the pNHMM2 continued to beable to reduce the pH versus the other hydrogels. On the other hand,after gelation and placement in fresh PBS, followed by 24 hstabilization, the pH was lower in and on the surface of pNHMM2 hydrogelcompared to pNHMM0, pNHMM0.5 and pNHMM1, as indicated by a pH-sensitivedye (LysoSensor, whose emission intensity ratio between 540 nm and 440nm increases as pH decreases). In addition, the pH of the interior ofthe pNHMM2 hydrogel was lower than the pH on the surface. No significantdifferences were found among the other 3 polymer types (data notavailable for pNHMM5 and pNHMM10 due to rapid degradation). Thedegradation rate increased significantly as MAA was added in increasingproportion into the polymer backbone, as shown by the weight lossprofile of the hydrogels in PBS, FIG. 8, top panel. Without MAA, pNHMM0needed over 5 mo to lose 50% weight in PBS, and the same loss requiredabout 2 mo, 1 mo, 1 wk and 1 d for MAA containing copolymers with theMAA feed ratio at 0.5% (pNHMM0.5), 1% (pNHMM1), 2% (pNHMM2), 5% (pNHMM5)and 10% (pNHMM10), as shown in FIG. 3b . The temporal weight lossprofiles of the pNHMMj hydrogels share a similar shape, which beginswith a slow weight loss stage, followed by an abrupt decrease inremaining weight. Furthermore, when incubated in PBS solution at pH 9.5,the abrupt weight loss for MAA10 was postponed for 1 d (FIG. 26). SincepNHMM5 and pNHMM10 hydrogels were considered to degrade too quickly forpotential in vivo applications, hydrogels with less MAA were selected ascandidates for further evaluation.

The transition temperature of MAA hydrogels in PBS did not shiftsignificantly as the MAA content in the polymer backbone was increased.As shown in FIG. 9, top panel, the rapid increase in shear modulus G0 ofpNHMM0, pNHMM0.5, pNHMM1 and pNHMM2 occurred within 17.5±2.5° C., whichrepresents the sol-gel transition as temperature rises. Accompanyingthis mechanical transition the copolymers were observed to becomeoptically opaque and form white gels. The abrupt increase in shearmodulus or viscosity required only a few sec at 37° C., hence whenpNHMMj hydrogels were injected into PBS at 37° C., the response wasrapid. As shown in FIG. 9, bottom panel, solution pH did notsignificantly affect the transition temperature of pNHMMj hydrogels,especially in the weak acidic range. The Young's modulus of pNHMMjhydrogels were all above 200 kPa and did not vary significantly withincreasing MAA content (FIG. 10), indicating that the degradation orabsorption rate for the hydrogels was decoupled from their initialmechanical stiffness in tension when the MAA content was low.

As shown in FIG. 12, panel round single cells were evenly distributedwhen rSMCs were encapsulated within the hydrogels. Using trypan blue toquantify viability, one day after encapsulation, over 85% cells werealive in pNHMM0, pNHMM0.5, pNHMM1 gels, while less than 70% cellssurvived inpNHMM2 (FIG. 12). Viabilitywas —85% in pNHMM0, pNHMM0.5,pNHMM1 gels 4 d after encapsulation, while viability was less than 60%in pNHMM2 (FIG. 11). Qualitatively assessing cells stained with aviability marker at the time of encapsulation showed that after 1 and 7d culture, viable cells remained spread throughout the hydrogels (FIG.11). Cells isolated at 7 d and cultured on TCPS for 7 d qualitativelyshowed an ability to proliferate (FIG. 27). The cytotoxicity of thedegradation products of pNHMMj hydrogels were also tested. Live/deadstaining showed that rSMC proliferation was not impeded by any of thecopolymer degradation products: overall metabolic activity of thecultures during the 7 d increased, suggesting cell proliferation, andfew dead cells were observed after recovery and culture on TCPS (FIG.13A). Slightly different from the findings from the encapsulationexperiment, degradation products of the pNHMM2 hydrogel did not showhigher cytotoxicity compared to its counterpart pNHMMj hydrogels (FIG.13B).

In addition, smooth muscle cell proliferation was not significantlyaffected by the hydrogel degradation products. FIG. 14 shows the grossappearance of a rat heart injected with MAPLA Gel, which was similar tothose injected with hydrogels with different MAA content.

In vivo degradation and absorption of hydrogels was tested in rathindlimb muscles. Hydrogels solidified immediately upon being injectedinto the muscles and formed distinct volumes, displacing the muscletissue, as shown in the left columns of FIG. 15. Fluorescent imagingrevealed that the hydrogels formed well-defined regions, with minimaldiffusion into the tissue. This effect was also demonstrated onimmunohistochemical staining sections obtained from tissue immediatelyafter injection (data not shown). After 21 d, the boundaries betweeninjected hydrogels and the leg muscles were more diffuse compared to thetime right after injection, particularly for pNHMM1 and pNHMM2, as shownin the right columns of FIG. 15. Fluorescent signals could still beobserved, showing that the hydrogels were not completely absorbed.However, the fluorescence intensities were weaker in all four groupscompared to the initial status. Unabsorbed hydrogels could be identifiedin immunohistochemical sections in all four groups as irregularly shapedregions encompassed by dense cell populations (FIG. 16, panel a). Formost of the polymers these regions of injection had minimal or nofluorescence, while for the pNHMM0 group this region was seen to havehigher green fluorescence in large, continuous areas. The dense cellpopulations surrounding the polymers were CD68 positive and identifiedas macrophages (FIG. 16, panels a and b). For pNHMM0, pNHMM0.5 andpNHMM1 hydrogels, macrophages encompassed the materials, however, didnot appear to infiltrate into the polymer. For the pNHMM2 group,macrophages infiltrated into the injection site, separating theremaining hydrogel volumes. Similar observation could be made on H&Estained sections. Macrophages gathered and formed capsules aroundinjected hydrogels as shown in FIG. 16, panel c. Macrophages can befound in some regions of unabsorbed pNHMM2 hydrogel, migrating from theborder zone. This was not observed with the other 3 polymer types.

Discussion

Different strategies have been reported to modulate the degradation ofthermally responsive hydrogels. For poly(ethyleneglycol)-poly(ε-caprolactone)-poly(ethylene glycol) (PEG-PCL-PEG),poly(ethylene glycol)-poly(lactide-co-glycolide)-poly(ethylene glycol)(PEG-PLGA-PEG) and similar block copolymer systems, changing themolecular weight of individual blocks and the molecular weight ratiobetween hydrophobic and hydrophilic blocks results in changes indegradation rate. For thermally responsive systems that employcrosslinking, adjusting the crosslinking density is effective. Inpoly(N-isopropylacrylamide) based hydrogels, the content of hydrolyticpendant groups has been shown to have a significant influence on bothhydrogel degradation rate and thermal transition behaviors. In othercases, introducing enzyme sensitive cleavage sites has proven aneffective mechanism.

In this study, pNHMMj hydrogels were synthesized and displayed tunabledegradation behavior in vitro and in vivo, supporting the designhypothesis that the addition of MAA as an autocatalyst for MAPLAhydrolysis would provide a method to accelerate the hydrogeldissolution. Manipulation of MAA content over a relatively small rangeresulted in widely varying degradation behavior, with 50% solubilizationoccurring in time frames from days to months, wider than what has beenpreviously reported in the literature. The lower pH of the supernatantimmediately after gelation and after a series of PBS/medium changes andstabilization steps indicated that the proton concentrations were higherin the environment for pNHMMj hydrogels with higher MAA content, asexpected. This result also corresponds with the pH being lower insideand on the surface of pNHMM2 compared to other pNHMMj hydrogels havingless MAA content, since greater MAA content would be the source of extraprotons which decreased the supernatant pH. In addition, the pH appearedto be lower inside pNHMM2 compared to the surface of the hydrogel,suggesting diffusion-driven depletion of protons towards the surface ofthe hydrogel. The four pNHMMj hydrogels studied all experienced anabrupt mass loss at different times, indicating a rapid increase inhydrolysis rate and polymer hydrophilicity. This can be explained withthe hindered diffusion of protons. As the proton concentration slowlybuilds up in the hydrogels, the autocatalytic effect would becomestronger, leading to faster accumulation of cleaved lactic acidmolecules and MAA residues, which in turn would contribute to a higherproton concentration. The fact that degradation of pNHMM10 (with highestMAA content among pNHMMj hydrogels) was slowed in weak basic buffersupports the proposed autocatalytic mechanism.

Hydrogel degradation was largely decoupled from both stiffness andthermal transition behavior (FIG. 9, bottom panel), probably because theMAA content was too low to affect the dominance of NIPAAm and highlyhydrophobic MAPLA effects on thermal transition, and increases inelectrostatic repulsion between MAA residues was not consequential.Theoretically, the ionizable MAA residues should increase pNHMMjcopolymer hydrophilicity and pH sensitivity. Peppas et al.systematically studied the thermal sensitivity and pH sensitivity ofpoly(NIPAAm-co-MAA) hydrogels. They found that the thermal transitiontemperature increased from 32° C. to 34.5° C. as MAA in thepoly(NIPAAm-co-MAA) hydrogels increased from 0 to 12 mol %; on the otherhand, the swelling ratio of the poly(NIPAAm-co-MAA) hydrogels increasedsignificantly between pH 5.3 and pH 5.7. The pH sensitivity was alsoobserved in a poly(NIPAAm-co-MAA) interpenetrating polymeric network. Inthe current study, the examined pH range (6.4-7.4) was narrow,corresponding to infarcted cardiac muscle as a potential applicationarea. The pH range studied fell above the pKa of poly(methacrylic acid)as its effect on thermal transition behavior was studied for pNHMMjhydrogels containing small amounts of MAA. The pH sensitivity was notsignificant in the examined range and phase change and mechanicalbehavior that were attractive for minimally invasive delivery could bemaintained, while the degradation rate could independently be optimizedfor a given application.

As expected, the pNHMM2 hydrogel showed faster in vivo degradation anddisappearance from the site of injection compared to the other threehydrogels studied. However, the in vivo degradation and absorption ofhydrogels qualitatively appeared to be substantially slower than invitro. Day 21 was chosen as an endpoint for the in vivo studies based onthe in vitro weight loss curve, expecting that almost all of the pNHMM2hydrogel and about 50% of the mass of the pNHMM1 hydrogel would havebeen lost. Furthermore, hydrogel mass loss was expected to potentiallybe faster in vivo if local enzymatic activity contributed to thedegradation process. The explanation for the apparently slower hydrogelloss in vivo could be that the catalyzing effect of MAA content wassomewhat weakened in vivo as the tissue served as a more effectivebuffering system relative to the PBS in vitro.

The fluorescence of the remaining injected hydrogels was diminished forthe faster degrading hydrogels with higher MAA content (FIG. 16). Thisresult was in accordance with what was observed in vitro. It is knownthat fluorescein changes from a lactone form to neutral and cation formsunder low pH, and its emission intensity and quantum yield dropsignificantly at ˜490 nm. After 21 d in PBS, the fluorescence intensityof the labeled hydrogel interior dropped significantly for pNHMM0.5 andpNHMM1 versus that measured immediately after gelation, whereas thefluorescence intensity of pNHMM0 remained unchanged (FIG. 28). Thisresult would explain the relatively strong green fluorescence only beingobserved for the pNHMM0 hydrogel after 21 d in vivo (FIG. 16),suggesting that the hydrogel interiors were more acidic for the otherthree studied hydrogels.

The foreign body response induced by the pNHMMj hydrogels 21 d postinjection was characterized by local macrophage recruitment, as istypical of the foreign body response. The macrophages mainly distributedaround the hydrogel mass, forming an encompassing cell layer. Fewforeign body giant cells could be observed. For the slow degradingpNHMM0, pNHMM0.5 and pNHMM1 hydrogels, the macrophages did notinfiltrate or separate the hydrogel mass, as opposed to the findingswith pNHMM2, where the hydrogel was fragmented and the pieces surroundedby macrophages. This suggested that the in vivo absorption of pNHMMjhydrogels is not simply characterized by surface mass loss, but thatdegradation and solubilization occurs across the hydrogel withfragmentation. This fragmentation was observed in all cases in vitro,with fragmentation not occurring until the rapid phase of mass lossbegan.

The principle of manipulating acid autocatalysis may apply forcontrolling degradation rates for other thermally responsive hydrogels,and more broadly, biomaterials whose degradation is based on polyesterhydrolysis. Ara et al. blended calcium compounds with different aciditywith PLGA and found that the most basic calcium carbonate was mosteffective in delaying the degradation, while the most acidic calciumdihydrogenphosphate was least effective. Similar buffering effects onautocatalysis were also found by Zhang et al., where basic magnesiumcarbonate and magnesium hydroxide significantly decreased thedegradation rate of PLGA films while neutral sodium chloride and zincsulfate did not show comparable influence. Wu et al. have shownexperimental results suggesting that pore morphology affects thediffusion of acidic degradation products from PLGA scaffolds, thusdirectly impacting the extent of acid catalyzed hydrolysis anddegradation rate. For pNHMM hydrogels, the mass loss curves werecharacterized by rapid degradation following early periods of slowerchange, similar to the behavior of several other polyester materials.The autocatalytic mechanism might readily be applied to the design andsynthesis for current material systems to introduce a tunable means ofincreasing degradation. For example, it is reasonable to assume thatadding a small portion of pendant carboxyl groups onto PEG-PCL-PEG,PCL-PEG-PCL, or PEG-PLGA-PEG hydrogels could accelerate degradation,although such a design might also be dependent upon the overallhydrophilicity of the modified polymer system. Of note, a syntheticroute to fabricate PCL-PEG-PCL copolymer bearing pendant carboxyl groupson the PCL blocks has been reported by Lavasanifar et al. Thedegradation rate of thermally responsive hydrogels has been studied andutilized to control the release of pharmaceutical agents and proteins.The potential of thermally responsive hydrogels as cell carriers hasbeen demonstrated as well. Given that hydrogel degradabilitysignificantly influences cell proliferation, migration anddifferentiation, along with the therapeutic outcomes of injectiontherapies, adding acid to tune the degradability of thermally sensitivehydrogels is attractive.

One key advantage of the strategy employed here is that relatively smallamounts of catalyst (acid) did not significantly influence the materialproperties of interest. However, there are potential limitations anddrawbacks of this strategy. First, for the acid to play its role, acertain level of diffusion and access of protons to the ester bonds isneeded, which in turn is dependent on the water content in the hydrogel.The accelerating effect may not be as readily achieved for morehydrophobic materials, particularly in the early stages of degradation.Also, for injectable hydrogels, the catalyzing effect may vary acrossinjection sites under the influence of buffering capacity. Second, anacidic microenvironment would be generated for faster degradation totake place, which would also lower the pH for the cells andcorresponding physiological events in the vicinity of the implant. Thiswas not explicitly studied in this report, but the histological resultsdid not suggest an obvious effect for the healthy skeletal muscle bedexamined

Conclusion

A series of NIPAAm based thermally responsive polymers (pNHMMj) weresynthesized, which contained varying MAA content. By releasing protonsthat catalyzed ester hydrolysis, MAA accelerated the degradation andadsorption of the hydrogels both in vitro and in vivo, and this effectwas positively correlated to MAA content in the copolymer. The thermaltransition behavior and mechanical strength did not significantly varyamong the pNHMMj hydrogels. On the other hand, the pNHMMj hydrogelsshowed low cytotoxicity and appeared to induce a mild foreign bodyresponse. The introduced autocatalysis strategy of tuning thedegradation of thermally responsive hydrogels where degradation orsolubilization is determined by their polyester components might beapplied to other tissue engineering and regenerative medicineapplications where appropriate biomaterial degradation behavior isneeded.

Example 6 Synthesis of poly(NIPAAm-co-VP-co-MAPLA) Hydrogels

Poly(NIPAAm-co-VP-co-MAPLA) copolymers were synthesized from NIPAAm, VPand MAPLA by free radical polymerization. The feed ratios of NIPAAm, VPand MAPLA were 80/j/(20-j), where j=10, 15 (FIG. 17). Monomers (0.066mol) were dissolved in 180 mL of 1,4-dioxane containing 0.23 g BPO. Thepolymerization was carried out at 70° C. for 24 h under an argonatmosphere. The copolymer was precipitated in hexane and furtherpurified by precipitation from THF into diethyl ether and vacuum-dried,with yields of ˜80%.

¹H spectra of poly(NIPAAm-co-VP-co-MAPLA) copolymers were recorded witha 600 MHz BRUKER spectrometer using CD3Cl as solvent (FIG. 17).Molecular weight of the copolymers was determined by gel permeationchromatography (GPC, Waters Breeze System, Waters 1515 HPLC PumpWaters2414 differential refractometer). The copolymers were dissolved in THFat a concentration of 1 mg/mL and the GPC analysis were performed at 35°C. A poly(methyl methacrylate) standard kit (Fluka, ReadyCal Set Mp500-2700000) was used for molecular weight-elution volume calibration.

Results: Poly(NIPAAm-co-VP-co-MAPLA) copolymers with differentcomposition were successfully synthesized. The actual compositions wereslightly different from the feed ratio (FIG. 17).

Rheology studies were conducted on a TA Instrument rheometer (AR2000) toobserve the mechanical property change of the hydrogels during thetemperature induced sol-gel transition. The polymer solutions (15 wt %in PBS) were placed between two parallel plates. The heating rate was 5°C./min, the shear storage modulus G′ were collected as a function oftemperature at a fixed strain of 2% and a frequency of 1 Hz.

To measure the transition time of hydrogels in 37° C. air, 150 uL ofeach type of hydrogel was added to pre-cooled a 96-well plate and placedin a plate reader which was pre-warmed to and set at 37° C. Absorbanceat 490 nm was recorded for 15 min.

Results: VP10 and VP15 have transition temperature of 19° C. and 26° C.,respectively (FIG. 18). VP15 has higher hydrophilicity as evidenced byhigher transition temperature and longer transition time.

Relevance: It shows our capability to modulate hydrogel hydrophilicityby changing polymer composition under the described design and synthesisroute.

Hydrogel degradation was quantified by mass loss measurements. Hydrogelswith known initial dry masses (˜ 60 mg) were immersed into 6 mL of PBS(replaced weekly to maintain a constant pH value of 7) at 37° C. Atpredefined time points over a 10 weeks period the hydrogel (n=3 each)were lyophilized and the relative mass loss was recorded.

Hydrogel discs after 24 h equilibrium in 37° C. PBS were weighed (wateron gel surface removed) and thoroughly dried in oven. Water absorptionwas calculated as (wet weight-dry weight)/wet weight×100%.

Results: Being more hydrophilic, VP15 absorbs more water and degradesfaster in PBS than VP10 (FIG. 19). 50% weight loss in PBS takes 5 weeksand 7 weeks for VP15 and VP10, respectively.

Relevance: Similar with other hydrogels disclosed,poly(NIPAAm-co-VP-co-MAPLA) hydrogels can provide mechanical support oninjection sites for a relatively long time.

To measure the mechanical properties of the hydrogels, samples wereincubated in a 37° C. water bath for 24 h to reach a stable watercontent, and then the solid hydrogels were cut into round discs with 3mm diameter and 3 mm thickness, and then loaded in a water bathequilibrated to 37° C. An ElectroForce 3200 Series II (Bose, Minnesota,US) equipped with a 2.5 N load cell was utilized to record thecompressive stress-strain curve immediately after the samples were takenout of the water bath.

Results: Being more hydrophilic, VP15 has a weaker mechanical strengththan VP10 (FIG. 20).

Relevance: The compressive modulus of VP15 is within the range of softtissue modulus, meaning it is strong enough as a bulking agent. VP10should work better in terms of mechanical strength.

The cytotoxicity of the degradation products of VP15 and VP10 hydrogelswas assessed by measuring the relative metabolic activity of ratmyoblasts (H9C2 cells) cultured in Dulbecco's modified Eagle medium(DMEM) (Gibco, Life technologies) with 10% fetal bovine serum (FBS), 1%penicillin/streptomycin, and supplemented at 10% with hydrogeldegradation solution. The hydrogel degradation solution was preparedfrom incubation of the hydrogel in PBS. Culture medium with PBS added at10% was used as a negative control. H9C2 cells were seeded at an initialdensity of 30,000/cm2 and their metabolic activity was measured (n=6each) using an MTS assay kit (Promega CellTiter 96 Cell ProliferationAssay). To qualitatively verify the results of the above test, cellswere also observed under fluorescence microscopy after live/deadstaining with a Promokine Live/Dead Cell Staining Kit.

Results: H9C2 cells stayed alive and proliferate normally under theinfluence of hydrogel degradation products (FIGS. 21 and 22).

Relevance: The degradation products of VP15 and VP10 have negligiblecytotoxicity, showing their safety and biomedical materials.

Adult female Lewis rats weighing 160-210 g were utilized in a protocolthat followed the National Institutes of Health guidelines for animalcare and that was approved by the University of Pittsburgh'sInstitutional Animal Care and Use Committee. Anesthesia was induced with3.0% isoflurane inhalation with 100% oxygen followed by 1.5-2%isoflurane with 100% oxygen during procedure. Dermatotomy was performedto expose the inner thigh muscles on both legs. Single injections of200-250 μL of VP15 or VP10 hydrogel were made approximately 3 mm deep inthe muscle bed. After 28 d, rats were sacrificed and the inner thighmuscles were excised and fixed in 10% formaldehyde for 3 d beforeembedding. H&E staining was performed (FIG. 23).

Results: Typical foreign body response was observed featuring recruitedmacrophages. However, the morphology of VP10 and VP15 28 days afterinjection is different. VP10 formed a solid mass in tissue while VP15diffused better and integrated with the muscle (FIG. 23).

Relevance: This difference shows VP10 and VP15 may have differentreleasing effects while carrying bioactive agents.

40 or 10 μL of adeno-associated virus (AAV) with packed GFP was mixedwith VP15 hydrogel before gelation at 37° C. Hydrogel was then immersedin 37° C. DMEM. Media with released AAV particles were collected atdifferent time points and new media was added to allow further release.The media with released AAV particles was used to infect pre-seededHEK293 cells for 48 h. Cells were observed with a fluorescentmicroscope.

Results: HEK293 cells were infected with AAV particles released fromhydrogel. The AAV particles released 14 days after gelation still hassome level of activity (FIG. 24).

Relevance: It is proved that slow release of bioactive agents can beachieved with poly(NIPAAm-co-VP-co-MAPLA) hydrogels while maintainingsome level of activity, which shows the potential ofpoly(NIPAAm-co-VP-co-MAPLA) hydrogels as mechanical support+carrier forbioactive agents.

The present invention has been described with reference to certainexemplary embodiments, dispersible compositions and uses thereof.However, it will be recognized by those of ordinary skill in the artthat various substitutions, modifications or combinations of any of theexemplary embodiments may be made without departing from the spirit andscope of the invention. Thus, the invention is not limited by thedescription of the exemplary embodiments, but rather by the appendedclaims as originally filed.

1. A composition, comprising a copolymer having a lower critical solution temperature (LCST) of less than 37° C., comprising N-alkyl acrylamide residues in which the alkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl; hydroxyethylmethacrylate (HEMA) residues; methacrylate-polyactide (MAPLA) macromer residues; and methacrylic acid (MAA) residues, and having a degradation rate of less than 200 days in vivo.
 2. The composition of claim 1, prepared by radical polymerization of a mixture of NIPAAm residues; HEMA residues, MAPLA macromer residues; and MAA residues.
 3. The composition of claim 1, wherein the N-alkyl acrylamide is NIPAAm and the copolymer is prepared by radical with a feed ratio of NIPAAm:HEMA:MAPLA:MAA in the range of 75-85:3-14.5:5-10:0.5-2.
 4. The composition of claim 1, wherein the N-alkyl acrylamide is NIPAAm and the copolymer is prepared by polymerization with a feed ratio of NIPAAm:HEMA:MAPLA:MAA of 80:9.5:10:0.5, 80:9:10:1, 80:8:10:2, or 80:5:10:5.
 5. The composition of claim 1, wherein the N-alkyl acrylamide is NIPAAm and the copolymer comprises NIPAAm:HEMA:MAPLA:MAA in a ratio having a range of 71.5-92.5:2.5-16:4.5-11:0.5-2.5, or 72-88:3-15:4.5-11:0.5-2.
 6. The composition of claim 1, wherein the copolymer has a degradation rate of less than 90 days in vivo.
 7. A composition, comprising a copolymer having an LCST of less than 37° C., comprising an N-alkyl acrylamide residue in which the alkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl; acrylic acid (AAc); a hydroxyethylmethacrylate-poly(trimethylene carbonate) (HEMAPTMC) macromer; and a methacrylic acid (MAA) monomer residue, and having a degradation rate of less than 200 days in vivo.
 8. The composition of claim 7, wherein the N-alkyl acrylamide is NIPAAm and the copolymer is prepared by polymerization with a feed ratio of NIPAAm:AAc:HEMAPTMC:MAA in the range of 75-85:3-14.5:5-10:0.5-2.
 9. The composition of claim 7, wherein the N-alkyl acrylamide is NIPAAm and the copolymer is prepared by polymerization with a feed ratio of NIPAAm:AAc:HEMAPTMC:MAA of 80:9.5:10:0.5, 80:9:10:1, 80:8:10:2, or 80:5:10:5.
 10. The composition of claim 7, wherein N-alkyl acrylamide is NIPAAm and the copolymer comprises NIPAAm:AAc:HEMAPTMC:MAA in a ratio having a range 71.5-94:2.5-16:4.5-11:0.5-2.5, or 72-88:3-15:4.5-11:0.5-2.
 11. A composition, comprising a copolymer having a lower critical solution temperature (LCST) of less than 37° C., comprising N-alkyl acrylamide residues in which the alkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl; N-vinylpyrrolidone (VP) residues; and methacrylate-polyactide (MAPLA) macromer residues; and having a degradation rate of less than 200 days in vivo.
 12. The composition of claim 11, prepared by radical polymerization of a mixture of NIPAAm residues; VP residues, and MAPLA macromer residues.
 13. The composition of claim 11, wherein the N-alkyl acrylamide is NIPAAm and the copolymer is prepared by polymerization with a feed ratio of NIPAAm:VP:MAPLA in the range of 70-90:5-20:5-20.
 14. The composition of claim 11, wherein the N-alkyl acrylamide is NIPAAm and the copolymer is prepared by polymerization with a feed ratio of NIPAAm:VP:MAPLA of 80:10:10 or 80:15:5.
 15. The composition of claim 11, wherein the N-alkyl acrylamide is NIPAAm and the copolymer comprises NIPAAm:VP:MAPLA in a ratio having a range of 63-93:4-22:3-22, or 68-92.5:4.5-21:3-21.
 16. The composition of claim 11, wherein the copolymer has a degradation rate of less than 90 days in vivo.
 17. A method of treating a myocardial defect in a patient comprising injecting a composition comprising a copolymer having a lower critical solution temperature (LCST) of less than 37° C. comprising: a. NN-alkyl acrylamide residues in which the alkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl; hydroxyethylmethacrylate (HEMA) residues; methacrylate-polyactide (MAPLA) macromer residues; and methacrylic acid (MAA) residues, and having a degradation rate of less than 200 days in vivo; b. N-alkyl acrylamide residues in which the alkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl; acrylic acid (AAC); hydroxyethylmethacrylate-poly(trimethylene carbonate) (HEMAPTMC) macromer residues; and methacrylic acid (MAA) monomer residues, and having a degradation rate of less than 200 days in vivo; or c. N-alkyl acrylamide residues in which the alkyl is one of methyl, ethyl, propyl, isopropyl and cyclopropyl; N-vinylpyrrolidone (VP) residues; and methacrylate-polyactide (MAPLA) macromer residues; and having a degradation rate of less than 200 days in vivo. into the area of a myocardial defect in the patient.
 18. (canceled)
 19. (canceled)
 20. (canceled)
 21. The method of claim 17, wherein the defect is an infarct or a traumatic injury, and the copolymer composition is injected into, in contact with and/or in tissue surrounding the infarct or traumatic injury.
 22. The method of claim 17, wherein the copolymer composition is injected into, in contact with, and/or in tissue surrounding the infarct or traumatic injury between 1 and 21 days after a myocardial infarction or myocardial injury.
 23. The method of claim 17, wherein the copolymer composition is injected into, in contact with, and/or in tissue surrounding the infarct or traumatic injury between 7 and 14 days after a myocardial infarction or myocardial injury.
 24. The method of claim 17, wherein the composition has a degradation rate of less than 90 days in vivo. 